Scanner Utilizing Beam Computed Tomography And Antiscatter Grid

ABSTRACT

A portable computed tomography (CT) system includes an O-shaped gantry defining an opening, an x-ray source operably coupled to the O-shaped gantry, and a flat panel detector (FPD) coupled to the O-shaped gantry and having a two-dimensional anti-scatter grid (2D ASG) coupled to a side of the FPD facing the opening. With the O-shaped gantry having the FPD, the object may be imaged in a first field of view (FOV) with the detector arranged in a centered geometry. Then, the detector may be arranged in an offset geometry, through-holes of the ASG may be aligned with x-ray emission paths of the x-ray source, and the object may be imaged in a second FOV with the detector arranged in the offset geometry.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation-in-part of U.S. patent applicationSer. No. 17/572,718 filed Jan. 11, 2022; which is a continuation of U.S.patent application Ser. No. 16/606,141 filed Oct. 17, 2019, and issuedas U.S. Pat. No. 11,224,389 on Jan. 18, 2022; which is a national stageof International Patent Application No. PCT/US2018/027660 filed Apr. 13,2018; which claims the benefit of and priority to U.S. ProvisionalPatent Application Nos. 62/486,113 filed Apr. 17, 2017, 62/573,021 filedOct. 16, 2017, and 62/576,265 filed Oct. 24, 2017; the entire contentsof each of which is incorporated herein by reference herein.

STATEMENT OF GOVERNMENTAL SUPPORT

This invention was made with government support under grant numbersR21CA198462 and R01CA245270 awarded by the National Cancer Institute ofthe National Institutes of Health. The government has certain rights inthe invention.

FIELD OF THE INVENTION

The present invention relates generally to X-ray detectors and moreparticularly to a system and a method for integrating an antiscattergrid with scintillators to significantly enhance the performance of aflat panel X-ray detector in portable computed tomography (CT) scanners.In particular, the performance of a flat panel X-ray detector in aportable CT scanner may be enhanced by integrating the flat paneldetector with 2D antiscatter grid.

BACKGROUND OF THE INVENTION

The purpose of a 2D antiscatter grid is to stop scattered x-raysreaching a flat panel detector, and improve the image quality. However,the design of the 2D antiscatter grid and the method to combine it withthe subcomponents of the x-ray detector is important to maximize theimage quality.

Moreover, a fraction of scattered x-rays can still pass through a 2Dgrid, and reach the detector. As a result, the image quality can bedeteriorated. Although, 2D grid's height (or grid ratio) can beincreased to reduce the transmission of scatter, such a grid design willlead to other technical and practical challenges, which deterioratesimage quality. Therefore, a correction algorithm is required to correctthe residual scatter intensity transmitted through the 2D grid. What isneeded is a 2D grid and methods that improve image quality andperformance of flat panel X-ray detector systems.

Additionally, a new function for the 2D grid is disclosed in thisinvention; a 2D grid generates high and low fluence regions in the x-raydetector due to the footprint of the 2D grid on the x-ray detector. Thismodulated fluence pattern can be used to solve the “pulse pile-up”problem in photon counting x-ray detectors.

Cone beam computed tomography (CBCT) is a compact 3D imaging technologythat has been utilized in a broad set of medical and industrial imagingapplications including dentistry, orthopedics, and interventionalradiology, among others. While CBCT provides cost efficient volumetricimaging capability, its main drawback is its relatively poor imagequality due to a large volume of x-ray illumination and scatteredradiation. In recent years, antiscatter grids have been a focus ofresearch to reduce X-ray scattering in CBCT imaging. However, thebenefits offered by commercially-available antiscatter grids do notsufficiently improve quality of soft tissue images for quantitativemeasurement. In clinical applications, poor soft tissue visualizationand reduced quantitative accuracy in CBCT impacts clinicaldecision-making. There exists a need to improve upon the antiscattergrid technology to increase CT number accuracy and contrast sensitivityin soft tissue imaging.

SUMMARY

One aspect of the disclosure provides a method of operating an x-rayimaging system. The method may include the step of arranging an x-raysource and a detector having an antiscatter grid (ASG) in a centeredgeometry wherein through-holes of the ASG are aligned with x-raysemission paths of the x-ray source. The method may include the step ofpositioning an object between the detector and the x-ray source. Themethod may include the step of imaging the object in a first field ofview (FOV) with the detector and the x-ray source arranged in thecentered geometry. The method may include the step of arranging thex-ray source and the detector in an offset geometry wherein thethrough-holes of the ASG are at least partially unaligned with thex-rays emission paths of the x-ray source. The method may include thestep of moving the detector in the offset geometry to realign thethrough-holes of the ASG with the x-ray emission paths of the x-raysource. The method may also include the step of imaging the object in atleast a second FOV with the detector and the x-ray source arranged inthe offset geometry.

Another aspect of the disclosure provides a portable computed tomography(CT) system. The system may include an O-shaped gantry defining anopening for a to-be-imaged object to be placed. The system may includean x-ray source operably coupled to the O-shaped gantry. The system mayinclude a flat panel detector (FPD) operably coupled to the O-shapedgantry. The FPD may include an x-ray absorbing sensor layer and adetector pixel array. The system may include an antiscatter grid (ASG)operably coupled to a side of the FPD facing the opening of the O-shapedgantry. The ASG may include a plurality of vertical walls definingopen-ended channels and formed of a radiation-absorbing material. Theopen-ended channels may be arranged in a geometric pattern pointedtoward the x-ray source to receive x-rays in an x-ray emission path fromthe x-ray source. A pitch between the vertical walls of the ASG may belarger than a pitch of the detector pixel array.

Yet another aspect of the disclosure provides an image data correctionmethod for an x-ray imaging system. The method may include the step ofreceiving, from a flat panel detector (FPD) of the x-ray imaging system,data representative of an x-ray image of at least a portion of anobject. The method may include the step of estimating a residual scatterintensity of x-ray radiation reaching the FPD from an x-ray source ofthe x-ray imaging system. Using the portable CT scanner according to thepresent technology, for example, correction of the residual scatterintensity may be accomplished, at least in part, by measuring signalintensity variations in the FPD or detector pixel array to facilitate animprovement in the quality of the x-ray image.

A portable CT system according to the present technology may use anO-shaped gantry, which allows for a robust implementation of anantiscatter grid (ASG) (e.g., a 2D ASG) and the disclosed residualscatter correction method. Correction of the residual scatter intensityaccording to the present technology may be accomplished, at least inpart, by measuring signal intensity variations in the FPD or detectorpixel array to facilitate an improvement in the quality of the x-rayimage An O-shaped gantry is mechanically more stable; a position of thex-ray source in relation to the detector does not change during gantryrotation. As a result of this mechanical stability, an O-shaped gantrymakes sure that a 2D-ASG's footprint (or shadow) does not change duringgantry rotation, which aids in the reliability of the disclosed scattercorrection method. Hence, image quality can be improved substantially.

The O-shaped gantry design according to the present technology alsoadvantageously provides sufficient mechanical support to x-ray tube anddetector, such that: a footprint of the ASG (e.g., 2D ASG) in the x-rayimage remains stationary throughout the gantry rotation, and theperformance of residual scatter intensity estimation method as disclosedherein may be improved.

The CT scanners and associated methods according to the presenttechnology may utilize flat panel detectors (FPDs) to enable imaging thewhole brain with one or two gantry rotations. This makes theelectromechanical system simpler and lighter. A lighter CT scanner ismore portable. Such a light gantry can also allow tilting of the gantryto accommodate patients who cannot lay flat. Known portable CT scannersneed a fast-rotating gantry that need to rotate many times (10-15 times)to be able to scan the whole brain. As a result, the gantry of such aknown system must be strong to handle the g-forces. In addition, theyrequire more x-ray output and a larger x-ray generator because they scana small portion of the brain per gantry rotation.

Known portable CT systems use conventional CT detectors that are customdesigned for each CT scanner. They are heavy and more expensive ascompared to the CT scanners of the present technology, which employ FPDsthat are lighter and cheaper than custom CT detectors. The CT scanneraccording to the present technology may be more suitable for CTperfusion and CT angio imaging than current portable CT scanners because3-5 cm thick brain tissue in the craniocaudal direction needs to beimaged repeatedly for perfusion CT imaging. Such a 3-5 cm thick field ofview can be achieved with the use of FPDs in the disclosed CT system,whereas existing CT systems may only cover 1-2 cm thick brain tissueduring perfusion CT imaging, which is suboptimal.

The disclosure also relates generally to X-ray detectors and moreparticularly to a system and methods for using an anti-scatter grid tosignificantly enhance the performance of flat panel X-ray detector. Inone embodiment, the invention comprises two major components: 1) anantiscatter grid design, which aims to reduce scatter intensity reachingthe flat panel detector (Schemes for integration of the antiscatter gridwith the flat panel detector are described. Additionally, a method tocorrect residual scatter intensity transmitted through the antiscattergrid are described). 2) a new antiscatter grid design and method weredescribed to solve the pulse pile-up problem in photon counting x-raydetectors.

This invention is described in preferred embodiments in the followingdescription with reference to the Figures, in which like numbersrepresent the same or similar elements. Reference throughout thisspecification to “one embodiment,” “an embodiment,” or similar languagemeans that a particular feature, structure, or characteristic describedin connection with the embodiment is included in at least one embodimentof the present invention. Thus, appearances of the phrases “in oneembodiment,” “in an embodiment,” and similar language throughout thisspecification may, but do not necessarily, all refer to the sameembodiment.

The described features, structures, or characteristics of the inventionmay be combined in any suitable manner in one or more embodiments. Inthe following description, numerous specific details are recited toprovide a thorough understanding of embodiments of the invention. Oneskilled in the relevant art will recognize, however, that the inventionmay be practiced without one or more of the specific details, or withother methods, components, materials, and so forth. In other instances,well-known structures, materials, or operations are not shown ordescribed in detail to avoid obscuring aspects of the invention.

The present invention relates generally to X-ray detectors and moreparticularly to a system and a method for using antiscatter grids tosignificantly enhance the performance of flat panel X-ray detector. Inone embodiment, the invention relates to a novel flat panel detector,which integrates a two dimensional (2D) antiscatter grid and a phosphorlayer with the detector's pixel array. In one embodiment, the twodimensional anti scatter grid of apertures separated by radio-opaque, orat least radiation-absorbing, septa, placed over, on, or in contactwith, the flat panel detector (FPD). In an example, a gap may existbetween the ASG and the FPD. In one embodiment, 3D printing is usedconstruct a 2D aperture array, such that each aperture may be preciselyaligned towards the x-ray focal spot. In one embodiment, lithographictechniques are used construct a 2D aperture array, such that eachaperture may be precisely aligned towards the x-ray focal spot. In oneembodiment, the presence of thin septa combined with the absence ofinter-septal spacers improves primary transmission, while the twodimensional aperture array provides efficient scatter rejectioncapability at levels not achievable with current scatter rejectiondevices.

In one embodiment, the invention relates to a signal adjustmentmechanism to counteract the shadow effect that the antiscatter grid mayhave on the detector. In one embodiment, the flat panel detector is beredesigned such that the 2D grid is directly placed on the phosphorlayer. In one embodiment, the normal flat panel detector continuousphosphor layer is replaced with a pixelated phosphor layer. In oneembodiment, the pixels in the phosphor are separated with reflectivesepta, preventing diffusion of visible light photons neighboringdetector pixels. In one embodiment, all or some of the walls of the 2Dgrid are aligned with the septa in the phosphor layer, minimizing the“inactive” area of the flat panel detector. It is believed that benefitcannot be achieved with the conventional approach where the grid ismounted on the top of the protective detector cover because the wallscannot be perfectly aligned with the septa of phosphor layer. It isbelieved that advantages of this hybrid design over existing flat paneldetectors include: Pixelated phosphor structure reduces cross-talkbetween detector pixels, improves spatial resolution; 2D grid providesbetter scatter absorption and improved primary x-ray transmission withrespect to conventional antiscatter grids. In return, it reduces noisein x-ray images. Integration of pixelated phosphor with the 2Dantiscatter grid reduces the percentage of “inactive” detector area,thus more primary x-rays will be detected by the detector.

In one embodiment the present invention contemplates an x-ray imagingdevice, comprising: a) an x-ray source; b) a two dimensional selectiveelectromagnetic radiation transmission grid comprising: i) a pluralityof vertical walls, wherein the walls are connected to each other in ageometric pattern and pointed towards the x-ray source; ii) a pluralityof open ended channels, wherein the open ended channels are defined bythe plurality of vertical walls; and c) a flat panel detector comprisinga phosphor layer and a detector pixel array, wherein the grid is incontact with the flat panel detector. In one embodiment, the verticalwalls comprise an essentially radiation-opaque, or at leastradiation-absorbing, material. In one embodiment, the vertical wallscomprise an at least partially radiation-opaque, or at leastradiation-absorbing, material. In one embodiment, the phosphor layer iscomposed of pixels, and pixels are separated with reflective walls. Inone embodiment, the vertical walls of the grid are aligned with thereflective walls of the phosphor layer. In one embodiment, the phosphorlayer is continuous, and the grid is placed directly on the phosphorlayer. In one embodiment, the separation (pitch) between the grid'svertical walls is larger than the pitch of the detector's pixel array,and vertical walls are not aligned with detector's pixel array. In oneembodiment, the grid's pitch and height may vary spatially across theflat panel detector. In one embodiment, there is a gap between the gridand phosphor layer. It is believed that this approach makes integrationof large area grids with large area flat panel detectors feasible. Inone embodiment, the flat panel detector comprises amorphous siliconpixel array. In one embodiment, the flat panel detector furthercomprises a complementary metal oxide semiconductor (CMOS) pixel array.In one embodiment, the geometric pattern is rectangular. In oneembodiment, geometric pattern is hexagonal. In one embodiment, the atleast one structural element comprises a plurality of radio-opaque, orat least radiation absorbing, sheets, such as metal sheets. In oneembodiment, the element extends at least between two of the plurality ofradio-opaque, or at least radiation-absorbing, sheets and on both sidesof at least one of the radio-opaque, or at least radiation absorbing,sheets. In one embodiment, the grid is a scatter measurement andcorrection grid, such as described in Example 7. In one embodiment, thedevice further comprises a correction algorithm. In one embodiment, thecorrection algorithm corrects a transmitted residual scatter intensity.

In one embodiment, the present invention contemplates an x-ray imagingdevice, comprising: a) an x-ray source; b) a two dimensional selectiveelectromagnetic radiation transmission grid comprising: i) a pluralityof septa, wherein the septa are connected to each other in a geometricpattern and pointed towards the x-ray source; ii) a plurality of openended channels, wherein the open ended channels are defined by theplurality of septa; and c) a flat panel detector configured underneaththe plurality of septa and comprising a plurality of photon countingpixels, wherein the grid is over, on, or in contact with, the flat paneldetector. In an example, a gap may exist between the ASG and the FPD. Inone embodiment, there is a gap between the grid and the flat paneldetector (or sensor). In one embodiment, the septa comprise anessentially radiation-opaque, or at least radiation-absorbing, material.In one embodiment, the flat panel detector comprises a Cadmium Telluridex-ray sensor. In one embodiment, the flat panel detector comprisesCadmium Zinc Telluride x-ray sensor. In one embodiment, the flat paneldetector comprises a Silicon x-ray sensor. In one embodiment, the flatpanel detector further comprises a complementary metal oxidesemiconductor (CMOS) pixel array. In one embodiment, the geometricpattern is rectangular. In one embodiment, the geometric pattern ishexagonal. In one embodiment, the at least one structural elementcomprises a plurality of radio-opaque, or at least radiation absorbing,sheets. In one embodiment, the sintered element extends at least betweentwo of the plurality of sheets and on both sides of at least one of thesheets. In one embodiment, the grid is a scatter measurement andcorrection grid. In one embodiment, the device further comprises acorrection algorithm. In one embodiment, the correction algorithmcorrects a transmitted residual scatter intensity. Correction of theresidual scatter intensity may be accomplished, at least in part, bymeasuring signal intensity variations in the FPD or detector pixel arrayto facilitate an improvement in the quality of the x-ray image. In oneembodiment, each radio-opaque, or at least radiation-absorbing, sheet inthe grid is composed of sections with different height and thicknesses.In one embodiment, the grid's shadow generates high and low x-rayfluence regions, or fluence modulation, incident on the flat paneldetector. It is believed that the fluence modulation pattern iscontrolled by changing the physical properties of the grid. In oneembodiment, the modulation pattern is used to correct pulse pile up inthe photon counting pixels.

In one embodiment, the present invention contemplates an x-ray imagingdevice, comprising: a) an x-ray source; b) a two dimensional selectiveelectromagnetic radiation transmission grid comprising: i) a pluralityof vertical walls, wherein the walls are connected to each other in ageometric pattern and pointed towards the x-ray source; ii) a pluralityof open ended channels, wherein the open ended channels are defined bythe plurality of vertical walls; and c) a flat panel detector comprisinga phosphor layer and a detector pixel array, wherein the grid is over,on, or in contact with, the flat panel detector. In an example, a gapmay exist between the ASG and the FPD. In one embodiment, the verticalwalls comprise an essentially radiation-opaque, or at leastradiation-absorbing, material. In one embodiment, the phosphor layer isdivided into pixels with reflective walls. In one embodiment, thevertical walls of the grid are aligned with the reflective walls of thephosphor layer. In one embodiment, the phosphor layer is continuous, andthe grid is placed directly on the phosphor layer. In one embodiment,there is a gap between the phosphor layer and the grid. In oneembodiment, the separation (pitch) between the grid's vertical walls islarger than the pitch of the detector's pixel array, and vertical wallsare not aligned with detector pixel array. In one embodiment, the grid'spitch and height may vary spatially across the flat panel detector. Inone embodiment, there is a gap between the grid and phosphor layer. Inone embodiment, the flat panel detector comprises amorphous siliconpixel array. In one embodiment, the flat panel detector furthercomprises a complementary metal oxide semiconductor (CMOS) pixel array.In one embodiment, the flat panel detector further comprises asolid-state x-ray sensor. In one embodiment, the geometric pattern isrectangular. In one embodiment, the geometric pattern is hexagonal. Inone embodiment, at least one structural element comprises a plurality ofradio-opaque, or at least radiation-absorbing, sheets. In oneembodiment, the structural element extends at least between two of theplurality of radio-opaque, or at least radiation-absorbing, sheets andon both sides of at least one of the radio-opaque, or at leastradiation-absorbing, sheets. In one embodiment, the grid is a scattermeasurement and correction grid. In one embodiment, the device furthercomprises a correction algorithm. In one embodiment, the correctionalgorithm corrects a transmitted residual scatter intensity.

In one embodiment, the present invention contemplates an x-ray imagingdevice, comprising: a) an x-ray source; b) a two dimensional selectiveelectromagnetic radiation transmission grid comprising: i) a pluralityof septa, wherein the septa are connected to each other in a geometricpattern and pointed towards the x-ray source; ii) a plurality of openended channels, wherein the open ended channels are defined by theplurality of septa; and c) a flat panel detector configured underneaththe plurality of septa and comprising a plurality of photon countingpixels, wherein the grid is over, on, or in contact with, the flat paneldetector. In an example, a gap may exist between the ASG and the FPD. Inone embodiment, the two dimensional grid further comprises at least onestructural element. In one embodiment, the septa comprises anessentially radiation-opaque, or at least radiation-absorbing, material.In one embodiment, the flat panel detector comprises silicon x-raysensor. In one embodiment, the flat panel detector comprises a cadmiumtelluride x-ray sensor. In one embodiment, the flat panel detectorcomprises a cadmium zinc telluride x-ray sensor. In one embodiment, theflat panel detector further comprises a complementary metal oxidesemiconductor (CMOS) pixel array. In one embodiment, the geometricpattern is rectangular. In one embodiment, the geometric pattern ishexagonal. In one embodiment, at least one structural element comprisesa plurality of radio-opaque, or at least radiation-absorbing, sheets. Inone embodiment, the radio-opaque, or at least radiation-absorbing, sheetin the grid is composed of sections with different height andthicknesses. In one embodiment, the grid's shadow generates high and lowx-ray fluence regions, or fluence modulation, incident on the flat paneldetector. In one embodiment, the modulation pattern is used to correctpulse pile up in the photon counting pixels. In one embodiment, thestructural element extends at least between two of the plurality ofradio-opaque, or at least radiation-absorbing, sheets and on both sidesof at least one of the radio-opaque, or at least radiation-absorbing,sheets. In one embodiment, the grid is a scatter measurement andcorrection grid. In one embodiment, the device further comprises acorrection algorithm. In one embodiment, the correction algorithmcorrects a transmitted residual scatter intensity. In one embodiment,the grid provides a fluence modulation on the flat panel detector. Inone embodiment, the grid has septa with uniform thickness. In oneembodiment, the grid comprises different thicknesses and heights insepta to control x-ray fluence modulation pattern. In one embodiment,the grid's septal shadows provide lower fluence.

In one embodiment, the present invention contemplates an x-ray imagingdevice, comprising: a) an x-ray source; b) a two dimensional selectiveelectromagnetic radiation transmission grid comprising: i) a pluralityof vertical walls, wherein the walls are connected to each other in ageometric pattern and pointed towards the x-ray source; ii) a pluralityof open ended channels, wherein the open ended channels are defined bythe plurality of vertical walls; and c) a flat panel detector comprisinga phosphor layer, wherein the grid is over, on, or in contact with, theflat panel detector. In an example, a gap may exist between the ASG andthe FPD.

Other objects, advantages, and novel features, and further scope ofapplicability of the present invention will be set forth in part in thedetailed description to follow, taken in conjunction with theaccompanying drawings, and in part will become apparent to those skilledin the art upon examination of the following, or may be learned bypractice of the invention. The objects and advantages of the inventionmay be realized and attained by means of the instrumentalities andcombinations particularly pointed out in the appended claims.

Definitions

To facilitate the understanding of this invention, a number of terms aredefined below. Terms defined herein have meanings as commonly understoodby a person of ordinary skill in the areas relevant to the presentinvention. Terms such as “a”, “an” and “the” are not intended to referto only a singular entity, but include the general class of which aspecific example may be used for illustration. The terminology herein isused to describe specific embodiments of the invention, but their usagedoes not delimit the invention, except as outlined in the claims.

As used herein, the term “3D printing” describes a manufacturing processfor metals or metal containing compounds to fabricate antiscatter grids.

As used herein, the term “sintered element” describes a material whichmay be used as a source material in a 3D printing process, such asdirect metal laser sintering. Selective laser sintering allows for alarge design freedom. Having a structural element that is built byselective laser sintering, the grid may be a highly complexthree-dimensional structure that is not easily achievable byconventional molding or milling techniques. Therein, the technology ofselective laser sintering, sometimes also known as direct metal lasersintering, is not any longer a prototype technology but becomes aproduction technology for the manufacturing of three-dimensional deviceswith demanding geometries.

As used herein, the term “essentially radiation-opaque material” is usedherein to describe a material, which does not allow for the transmissionof x-ray radiation. Examples of such materials are tungsten, tantalum,and lead. On the other hand, a “radiation-absorbing material” may absorbsome, but not all, x-rays. The ASG according to the present technologymay achieve its technically advantageous functions without being made ofa “radiation-opaque material” that fully absorbs the x-rays, but ratherin some embodiments can at least partially be made of a“radiation-absorbing material.”

As used herein, the term “metal oxide semiconductor (CMOS)” is usedherein to describe a technology for constructing integrated circuits.CMOS is also sometimes referred to as complementary-symmetrymetal-oxide-semiconductor (or COS-MOS). The words“complementary-symmetry” refer to the fact that the typical design stylewith CMOS uses complementary and symmetrical pairs of p-type and n-typemetal oxide semiconductor field effect transistors (MOSFETs) for logicfunctions. Two important characteristics of CMOS devices are high noiseimmunity and low static power consumption.

As used herein, the term “amorphous silicon (aSi) pixel array” is usedherein to describe a technology for constructing flat panel x-raydetectors. aSi is often used in construction of photodiodes and thinfilm transistor (TFT) arrays employed in flat panel x-ray detectors.

As used herein, the term “energy integrating” is used herein to describea type of x-ray detector, where only the accumulated x-ray energy in thex-ray detector is used to form the image. The number and energy ofindividual x-rays absorbed by the detector are not registered. An aSipixel array integrated with a phosphor layer is an example for an energyintegrating detector.

As used herein, the term “photon counting” is used herein to describe atype of x-ray detector, where energy of individual x-rays and the numberof x-rays interacting with the detector are recorded. Cadmium ZincTelluride (CZT) or Cadmium Telluride (CdTe) x-ray sensors integratedwith energy resolving readout electronics are examples for photoncounting x-ray detectors.

As used herein, the term “flat panel x-ray detector (FPD)” is usedherein to describe an “area” x-ray detector that can generate a twodimensional x-ray image. FPD can be an energy integrating or a photoncounting x-ray detector.

Abbreviations

ASG=antiscatter grid

2D ASG=two-dimensional antiscatter grid

FPD=flat panel detector

FDK=filtered backprojection

ART=adaptive radiation therapy

CT=computed tomography

CBCT=Cone beam computed tomography (also referred to as C-arm CT, conebeam volume CT, or flat panel CT)

MDCT=multi-detector CT

CTDI=CT dose index

CAX=projected location of the beam's central axis

SPR=Scatter-to-Primary ratio

SNR=signal-to-noise ratio

K_(SNR)=SNR improvement factor

T_(S)=scatter transmission fraction

T_(P)=primary transmission fraction

I_(S)=average scatter intensity

ROI=region of interest

BT=bow tie

DMLS=direct metal laser sintering

RT=radiation therapy

MTF=modulation transfer function

CT-to-ED=CT number to electron density

DVH=dose volume histograms

PTH=primary transmission histograms

MLEM=maximum likelihood expectation maximization

BRIEF DESCRIPTION OF THE DRAWINGS

The patent or application file contains at least one drawing executed incolor. Copies of this patent or patent application publication withcolor drawing(s) will be provided by the Office upon request and paymentof the necessary fee.

The accompanying figures, which are incorporated into and form a part ofthe specification, illustrate several embodiments of the presentinvention and, together with the description, serve to explain theprinciples of the invention. The figures are only for the purpose ofillustrating a preferred embodiment of the invention and are not to beconstrued as limiting the invention.

FIG. 1 shows an illustration of one embodiment of the two-dimensionalfocused antiscatter grid (2D ASG). Note the focusing property of the 2Darray: each grid element (i.e. through-hole) is aligned, or pointed,towards the x-ray source to account for the divergence of the x-raybeam.

FIG. 2 shows two pictures of a 2×20 cm2 2D antiscatter grid (ASG)module. Each grid hole has a unique slant, or angle, such that they arealigned towards the focal spot in half-fan CBCT geometry.

FIG. 3 shows (Top) Side view of the experimental setup. Red circlescorrespond to measurement points, and they are located at CAX, 10 cm,and 20 cm off-axis from CAX at detector plane. (Bottom) Beam eye view ofthe FPD. 2×40 cm2 wide “measurement area” of the FPD is shown in gray.Rest of the FPD was covered with Pb sheet.

FIG. 4A shows SPR values for 20 and 40 cm thick phantoms were plotted asa function of beam stop diameter. Linear fits were used to extrapolateSPR and T_(S) values to “0” beam stop diameter. Measurements wereperformed at 10 cm from CAX, and using Experiment Setup 2 in Table 1.

FIG. 4B shows T_(S) values for 20 and 40 cm thick phantoms were plottedas a function of beam stop diameter. Linear fits were used toextrapolate SPR and T_(S) values to “0” beam stop diameter. Measurementswere performed at 10 cm from CAX, and using Experiment Setup 2 in Table1.

FIG. 5A shows a section of the primary transmission map of 2D ASG, wherebright and dark regions indicate higher and lower primary transmissionvalues, respectively. The orange colored square indicates a 1.6×1.6 cm²ROI used for average primary transmission, T_(P), calculation.

FIG. 5B shows T_(P) as a function of ROI location along the long edge ofthe primary transmission map.

FIG. 5C shows cumulative primary transmission histograms of 1D and 2DASGs.

FIGS. 6A, 6B and 6C show SPR variation as a function of phantomthickness at 3 measurement points (CAX, 10 cm from CAX, and 20 cm fromCAX, respectively).

FIG. 7A shows the effect of air gap on SPR as a function of measurementlocations.

FIG. 7B shows the effect of BT filter on SPR as a function ofmeasurement locations.

FIGS. 8A, 8B and 8C show T_(S) as a function of phantom thickness at 3measurement points (CAX, 10 cm from CAX, and 20 cm from CAX,respectively).

FIG. 9A shows T_(S) as different air gaps were plotted at all 3measurement points.

FIG. 9B shows T_(S) measured with and without BT filter.

FIG. 10 shows SNR improvement factors, K_(SNR), were plotted as afunction of SPR without ASG. K_(SNR) values were calculated for allmeasurement points and experiment setups (See Table 1). Cubicpolynomials (red and blue lines) were fitted to better visualize trendsin K_(SNR).

FIG. 11 shows measured SPR with and without 2D ASG prototype versusthickness of slab phantom. At 20 cm phantom thickness, the 2D ASGsuppressed SPR from 3 to 0.1 at both 80 and 120 kVp [1].

FIG. 12 shows SNR improvement factor of one embodiment of the 2D ASG(solid line, circles) and a radiographic ASG with 1D lead septa (dottedline, triangles) [1-3]. Both ASGs have the same grid ratios (i.e. 12).SNR improvement factor below 1 at a given scatter to primary ratioindicates that ASG degrades the SNR.

FIG. 13 shows the average primary transmission fraction through oneembodiment of the 2D ASG as a function of ASG's grid pitch and septalthickness. The average primary transmission fraction of radiographicASGs is 60-70%, indicated by red rectangle.

FIG. 14A shows one embodiment of the 2D ASG built for FPD and CBCTapplications.

FIG. 14B shows the shadow, or footprint, of the 2D ASG appears as adarker grid pattern in the FPD image. X-ray intensity reaching the FPDis lower in the shadow, since 2D ASG's septa partially block primaryx-rays. If this intensity variation is not corrected, it will lead toimage artifacts, and makes the use of 2D ASG unfeasible.

FIG. 15 shows an example of the configuration of a current x-ray flatpanel detector and 1D anti-scatter grid.

FIG. 16 shows one embodiment of the current invention 2D grid placed onthe phosphor layer.

FIG. 17 shows the percent primary x-ray transmission fraction wascalculated as a function of grid's septal thickness (i.e. wallthickness) and grid pitch (i.e. grid's channel width) in the 2D ASGmodel. For a 100 micron thick wall and 3 mm wide channels, 85% of theprimary x-ray is transmitted through the grid.

FIG. 18 shows the percentage of detector pixels that receive at least50% of the nominal signal (PVH50) with the 2D grid. Without the 2D grid,all pixels will receive 100% of the nominal signal. When the 2D grid isin place, detector pixels in the shadow of the 2D grid will receivereduced signal. For example, for a grid channel width (pitch) of 3 mmand wall thickness of 100 microns (red curve), 95% of the pixels willstill receive at 50% of the nominal signal. Only 5% of the pixels willreceive less than 50% of the nominal signal. This 5% of the pixels areconsidered “dead” pixels. As the signal they receive is significantlyreduced due to the 2D grid's shadow. Markers indicate the calculatedvalues by the model, and lines are polynomial fits.

FIG. 19 shows an X-ray image with a 2D grid in place. Box 1 shows aregion underneath the 2D grid's footprint, and Box 2 indicates a regionat the center of a grid hole. The ratio of image intensity in Box 1 andBox 2 is a unique value. This intensity ratio changes as a function ofscatter content in the image. This ratio is measured without scatterpresent in the image (calibration data). Calibration data can later beused to estimate the scatter intensity in patient images. Such “boxpairs” can be created for other points underneath the grid's footprintand grid hole centers. As a result, scatter intensity can be estimatedfor any point in the image.

FIG. 20 shows the diffusion of x-rays with a focused 2D antiscatter gridwith a continuous scintillator and pixel array. This shows that therecan be a dip in signal due to shadow of the grid, depending on pixellocation.

FIG. 21 shows the diffusion of x-rays with a focused 2D antiscatter gridwith a continuous scintillator and pixel array. This shows that therecan be signal interference due to long-distance cross-talk.

FIG. 22 shows the diffusion of x-rays with a focused 2D antiscatter gridwith a pixelated phosphor layer and pixel array. This shows that signalcross-talk may be eliminated with this design.

FIG. 23 presents a representative x-ray fluence chart demonstrating the“pulse pile-up” problem.

Y axis: Counts recorded by the “photon counting” detector

X axis: True number of counts incident on the detector

Ideal detector response is the dashed black line

Colored lines are the actual detector responses for 3 differentdetectors.

In the region indicated by the red transparent rectangle, detectors cansignificantly “underestimate” the true x-ray counts.

FIGS. 24A and 24B presents the generation of an x-ray image by a 2-Dhexagonal grid.

FIG. 24A: A 2D antiscatter grid comprised of a series of parallelhexagonal holes (red center line).

FIG. 24B. An image intensity profile generated by x-ray intensitythrough the hexagonal grid of FIG. 24A is shown as an orange line.

FIG. 25 illustrates one embodiment of aft fluence modulation regions(yellow arrows) resulting from pixels within the septal shadow of a 2DASG grid.

FIG. 26 presents one embodiment of a photon counting detector.

FIG. 27 presents exemplary data of the range of maximum incident countrates in MDCT and CBCT and the corresponding output count rates.Incident count rates in MDCT (red region) is in the range of 10⁸-10⁹ persec/mm², whereas count rates in CBCT (green region) is 10⁷ per sec/mm²or less (each pixel is assumed be 1 mm² in the graph). [6]

FIGS. 28A and 28B present one embodiment of a 2D ASG. FIG. 28B shows amagnification of the red inset box in FIG. 28A, with a US quarter coinshown for dimensional purposes.

FIG. 29 presents exemplary data of a scattered radiation to primaryradiation ratio (SPR) as a function of slab phantom thickness. Note:Radiographic 1D ASG and 2D ASG have SPR ratios of 10 and 8,respectively. [17].

FIGS. 30A, 30B and 30C present representative images showing improvedCBCT image quality with 2D ASG. Images were acquired using a linacmounted CBCT system, and hence, the effect of gantry sag were included.Elliptical phantom mimics human torso dimensions (30 cm×38 cm).

FIG. 30A: Without an antiscatter grid.

FIG. 30B: With conventional 1D ASG.

FIG. 30C: With a 2D ASG prototype, presented in this invention.

FIGS. 31A, 31B and 31C present two septa embodiments configured toperform an aft fluence modulation method.

FIG. 31A: Standard constant septal thickness (e.g., without a footing).X-ray counts in pixel underneath the septal footprint is lower due topartial obstruction of fluence by septa. To reduce x-ray counts further,septal thickness should be increased.

FIG. 31B: An alternative solution is to use septa with a footing. Sincethe height of footing is 0-3 mm, it partially transmits x-rays.

FIG. 31C: Representative X-ray count profiles for septa configurationsshown in FIG. 31A and FIG. 31B.

FIG. 32 presents a diagram of a centered detector CBCT geometry for aCBCT system according to a known embodiment.

FIG. 33 presents a diagram of an offset detector CBCT geometry for aCBCT system according to a known embodiment.

FIG. 34A again presents the 2D ASG of FIG. 1 .

FIG. 34B presents a schematic diagram of an x-ray imaging systemaccording to some embodiments of the present technology.

FIG. 35 presents the system shown in FIG. 34B where the 2D ASG accordingto the present technology is shifted laterally.

FIG. 36 presents an x-ray imaging system 3600 according to someembodiments of the present technology.

FIG. 37A presents a schematic diagram providing additional details forthe shifting and tilting technique described above for the system ofFIG. 36 .

FIG. 37B presents a flowchart of a method to improve (or assure)alignment of an antiscatter grid when the detector is shifted withrespect to the x-ray source, according to some embodiments of thepresent technology.

FIGS. 38A-38D present an x-ray imaging system utilizing a flat paneldetector (FPD) with a 2D ASG as shown in FIG. 34A, according to someembodiments of the present technology.

FIGS. 39A-39F present a compact or portable CT scanner system presentsutilizing a flat panel detector (FPD) with a 2D ASG as shown in FIG.34A, according to some embodiments of the present technology.

FIG. 39G presents a method for processing x-ray images that may be usedwith the system of FIGS. 39A-39F, according to some embodiments of thepresent technology.

FIG. 39H presents CT scan images obtained using the system of FIGS.39A-39F as compared to images obtained using known CT scanners.

FIGS. 40A and 40B present images in comparison between an ultracompactCT (uCT) scanner with a 2D antiscatter grid according to someembodiments of the present technology (e.g., the system of FIGS.39A-39F) and a helical CT scanner.

FIGS. 41A and 41B present a comparison of images obtained using a CBCTsystem and processed without and with, respectively, the techniques ofthe method 3902 shown in FIG. 39G.

FIGS. 42A and 42B present a comparison of images obtained using thestandard CBCT scanner and ultracompact CT scanner, respectively,according to the present technology.

FIGS. 43A, 43B and 43C present simulated images in comparison todemonstrate the effect of CT gantry shape on gantry flex and 2Dantiscatter grid performance.

FIG. 44 presents a diagrammatic representation of a machine, in theexample form, of a computer system within which a set of instructions,for causing the machine to implement or otherwise perform any one ormore of the techniques and methodologies of the present technologydescribed herein, may be executed.

DETAILED DESCRIPTION

A portable CT system according to the present technology may use anO-shaped gantry, which allows for a robust implementation of anantiscatter grid (ASG) (e.g., a 2D ASG) and the disclosed residualscatter correction method. Correction of the residual scatter intensityaccording to the present technology may be accomplished, at least inpart, by measuring signal intensity variations in the FPD or detectorpixel array to facilitate an improvement in the quality of the x-rayimage An O-shaped gantry is mechanically more stable; a position of thex-ray source in relation to the detector does not change during gantryrotation. As a result of this mechanical stability, an O-shaped gantrymakes sure that a 2D-ASG's footprint (or shadow) does not change duringgantry rotation, which aids in the reliability of the disclosed scattercorrection method. Hence, image quality can be improved substantially.

The CT scanners and associated methods according to the presenttechnology may utilize flat panel detectors (FPDs) to enable imaging thewhole brain with one or two gantry rotations. This makes theelectromechanical system simpler and lighter. A lighter CT scanner ismore portable. Such a light gantry can also allow tilting of the gantryto accommodate patients who cannot lay flat. Known portable CT scannersneed a fast-rotating gantry that need to rotate many times (10-15 times)to be able to scan the whole brain. As a result, the gantry of such aknown system must be strong to handle the g-forces. In addition, theyrequire more x-ray output and a larger x-ray generator because they scana small portion of the brain per gantry rotation.

Known portable CT systems use conventional CT detectors that are customdesigned for each CT scanner. They are heavy and more expensive ascompared to the CT scanners of the present technology, which employ FPDsthat are lighter and cheaper than custom CT detectors. The CT scanneraccording to the present technology may be more suitable for CTperfusion and CT angio imaging than current portable CT scanners because3-5 cm thick brain tissue in the craniocaudal direction needs to beimaged repeatedly for perfusion CT imaging. Such a 3-5 cm thick field ofview can be achieved with the use of FPDs in the disclosed CT system,whereas existing CT systems may only cover 1-2 cm thick brain tissueduring perfusion CT imaging, which is suboptimal.

The disclosure also relates generally to X-ray detectors and moreparticularly to a system and a method for integrating an antiscattergrid with flat panel X-ray detectors to significantly enhance theperformance of flat panel X-ray detector. In addition, the count rateand energy resolution performance of a photon counting flat panel X-raydetector may be enhanced by using the disclosed antiscatter grid.

For a specific example, in recent years, the utility of in-room CBCTimaging has been extensively investigated in the context of noveltreatment strategies in radiation oncology. In the framework of adaptiveradiation therapy (ART), patient treatments can be monitored for tumorand normal tissue anatomic changes by utilizing CBCT images acquiredduring the course of treatment. The role of CBCT imaging can be furtherexpanded to modify treatment plans to account for anatomic changes, andhence, improve tumor control and reduce toxicity. Although much researchhas been done on evaluating the potential benefits of patient-specifictreatment monitoring and ART in general, the utilization of CBCT imagesin the context of ART did not translate into clinical practice due topoor quality of CBCT images. Currently, the utility of CBCT imaging islimited to in-room patient setup corrections only.

Scattered radiation is widely acknowledged to be a major cause of imagequality degradation in CBCT [4, 5]. Scatter leads to poor soft tissuevisualization, reduces CT number accuracy, and generates image artifactsin CBCT images. To mitigate the scatter problem, last decade has seenwidespread investigation of scatter suppression methods and devices.However, the desired improvement in CT number (Hounsfield Unit) accuracyand low contrast sensitivity has not been achieved.

To overcome the shortcomings of current scatter suppression methods, thepresent invention relates to a scatter rejection device: atwo-dimensional focused antiscatter grid (2D ASG), for FPDs and CBCTsystems. With respect to existing antiscatter grids (ASG), oneembodiment the present invention design has fundamentally differentproperties: it consists of a 2D array of apertures separated byradio-opaque, or at least radiation-absorbing, septa, placed over, on,or in contact with, the flat panel detector (FPD). In an example, a gapmay exist between the ASG and the FPD. In one embodiment, each aperturemay be 2-5 mm in width, and 20-50 mm in height, and its central axisoriented toward the x-ray focal spot to match the divergence of theimaging beam. Furthermore, aluminum or fiber inter-septal spacers (usedin radiographic ASGs) will be eliminated due to the mechanical strengthof the tungsten 2D array.

In one embodiment, DMLS enables construction of a 2D aperture array withseptal thicknesses down to 100 microns, and each aperture will beprecisely aligned towards the x-ray focal spot. The presence of thinsepta combined with the absence of inter-septal spacers may improveprimary transmission, while the 2D aperture array will provide efficientscatter rejection capability at levels not achievable with currentscatter rejection devices. Based on preliminary investigations, oneembodiment of the current invention 2D ASG reduces scatter-to-primaryratio (SPR) by a factor of 30, whereas SPR reduction factor withradiographic ASGs is limited to 5 to 10. Additionally, 2D ASG maytransmit 85% to 90% of the primary x-rays to the FPD in contrast to60-70% primary transmission through radiographic ASGs.

Given the high level of primary transmission and efficient scattersuppression capability of the 2D ASG, it was determined that oneembodiment of the current design enables significant enhancement in softtissue visualization, and permit high level CT number accuracy. Toassess this, several steps were taken:

Step 1 A range of 2D ASG prototypes with varying grid size and gridratios were designed and constructed. Scatter rejection and primarytransmission properties of the prototypes were characterized todetermine the optimal 2D ASG geometry.

Step 2. Evaluate the impact of 2D ASG on CBCT image quality. In acomprehensive set of experiments, the improvement in CBCT image qualitymetrics were evaluated with respect to the gold standard, multi-detectorCT (MDCT) images.

It is important to note that CBCT is utilized in other applications suchas vascular imaging [6], maxillofacial surgery [7], and spinalprocedures [8]. The results of this investigation will therefore applyto other CBCT systems that use fixed source-detector geometry.

significance

While CBCT is the most frequently used in-room 3D imaging modality forpatient setup corrections in radiation therapy (RT) [9], it additionallyprovides unique clinical information, thus far not exploited in theclinical practice of radiation therapy; CBCT images are acquiredperiodically during patient setups, often on a daily basis, capturing 3Danatomy of both normal tissue and tumors throughout treatment. It hasbeen widely suggested that CBCT images can be utilized in treatmentmonitoring and plan modifications to enhance therapeutic ratio, keyconcepts in the adaptive radiation therapy paradigm (ART) [10]. Briefly,ART aims to improve tumor control and reduce toxicity by monitoringchanges in normal tissues and tumors, and adapting treatment plans tosuch changes during the course of treatment.

The benefits of various ART approaches have been investigated fornumerous disease sites; for example, Yang et al. showed that adaptivetreatment replanning significantly improved quality of life in patientswith nasopharyngeal carcinoma [11]. Several studies indicated that ARTapproach can significantly reduce parotid dose and improve dose coverageof targets in Head and Neck cancer [12-16]. In Bladder cancer, Foroudiet al. concluded that ART can provide better normal tissue sparing ratiothan conventional approaches [17]. In cervix cancer, Tyagi et al. statedthat ART approach is needed to prevent suboptimal treatment coverage andexcessive toxicity [18].

The utility of CBCT in ART has been severely limited due to poor imagequality. In treatment response monitoring, assessment of changes inanatomy is challenging due to poor soft tissue visualization [18-20]. Atboth the treatment replanning and adaptation phases, accuracy ofstructure delineation and deformable image registration deteriorate[21-24], and CBCT-based dose calculations are not clinically acceptabledue to lack of CT number accuracy [24]. Due to poor image quality,clinical scope of CBCT has been limited to patient setup correctionssince its introduction to clinical practice.

Due to the CBCT image quality issues described above, the vast majorityof ART research is performed in academic centers using repeathigher-quality helical CT scans (rather than readily available in-roomCBCT images). This approach is not feasible for most RT clinics due topersonnel and resources needed to handle the increased workload.Moreover, some of the ART methods, such as online treatment adaptationsbefore treatment delivery [17], cannot be deployed without an in-roomimaging system such as CBCT. Thus if the image quality of CBCT were tobe improved sufficiently, this would represent a breakthrough inenabling wide-spread implementation of ART strategies in clinicalpractice. Although not all RT patients may benefit from the ARTapproach, high quality CBCT images are still essential for identifyingpotential patients for treatment adaptations [19, 20, 25, 26].

Scattered radiation is one of the leading causes of poor image qualityin CBCT [4, 5]. As each CBCT projection has a large field of view,significant amounts of scattered radiation originating from the patientreach the flat-panel detector (FPD), contaminating the image signal.Depending on the imaged anatomy and imaging geometry, the magnitude ofscatter may exceed primary x-ray intensity by a factor of 3 to 5 in CBCTprojections [4, 27-29]. Scattered radiation impacts 3 important aspectsof image quality: 1) Scatter deteriorates the contrast between softtissue structures. 2) Scatter greatly reduces accuracy of CT numbers,which is critical to accurate dose calculations in CBCT images. 3)Scatter induces image artifacts that impact both tissue visualizationand quantitative accuracy.

To improve CT number accuracy in CBCT, a wide range of scattercorrection methods have been investigated in the past decade, and suchmethods have been shown to improve CBCT image quality [29, 35-38].However, the improvements have been insufficient due to two fundamentalproblems: 1) All scatter correction methods use a model to estimate thescatter signal as a function of multiple model parameters. Assumptionsmade in the model lead to discrepancies between the estimated scatterand true scatter under real imaging conditions. 2) In scatter correctionmethods, “scatter-free” image signal is restored after detection ofscatter by the FPD. Thus, even an ideal correction method can onlyaccount for the bias in image signal amplitude due to scatter, andstochastic noise due to scatter cannot be removed from the image signal.Thus, soft tissue visualization cannot be improved due to stochasticnoise of scatter [5, 28, 29]. To physically suppress scatter,conventional radiographic ASGs (constructed from alternating layers oflead septa and spacer strips) were extensively investigated for CBCT[39-41]. While reduction in image artifacts and improvement in accuracyof CT numbers were demonstrated, these fell short of desired imagequality levels in CBCT [27, 42]. Due to relatively poor primarytransmission through the radiographic ASGs, only modest improvements inlow contrast resolution were achieved, and only under high scatterconditions [39, 43, 44].

An ideal solution would be to remove scatter before it is detected bythe FPD, using an efficient scatter suppression device, which alsomaintains high primary transmission efficiency. Thus, the presentinvention a two dimensional focused antiscatter grid (2D ASG) dedicatedfor CBCT systems was investigated and developed. The preliminaryassessment of the present invention indicates that the performance ofthe design may significantly close the gap in quantitative accuracybetween CBCT and MDCT, and improve low contrast resolution to levels notachieved by existing scatter suppression devices and correction methods.Fabrication of the advanced design is not feasible using standardmanufacturing techniques, and therefore new 3D printing technology maybe used to construct the current invention 2D ASG.

Innovation

The present invention antiscatter grid design has the promise tosignificantly improve CT number accuracy and low contrast resolution forCBCT systems.

The present invention grid architecture represents a fundamentaldeparture from existing antiscatter grid designs for FPDs. All currentantiscatter grids for FPDs are based on the 1D Potter-Bucky gridarchitecture [2, 45] (i.e. a 1D grid is formed by stacking alternatingstrips of lead and spacers made from aluminum or fiber). ThePotter-Bucky grid's inherent physical characteristics make it unsuitablefor construction of 2D anti-scatter grids.

Thus, in the present invention, a novel grid architecture is utilized(FIG. 1 ): The present invention device consists of 2D array of squarecross-sectioned grid elements separated by tungsten septa. The grid canbe attached to the protective cover of the FPD, or it can be directlyintegrated with the x-ray absorbing sensor layer. To account for x-raydivergence in cone beam geometry, each grid element is aligned, orfocused, towards the point x-ray source to maximize transmission ofprimary x-rays across the FPD. With respect to 1D Potter-Bucky grid, wereduced the grid pitch by more than an order of magnitude, and increasedthe septal thickness by a factor of 3 to 5 in the 2D grid (detailsdescribed in Step 1). A radio-opaque, or at least radiation-absorbing,material, such as tungsten, was utilized to provide mechanical strengthand higher radio-opacity, and eliminated inter-septal spacers. Thesedesign choices in the 2D ASG design were made to achieve both favorableimaging performance and ease of fabrication.

In the present invention, 2D ASG's septa do not have to be aligned withthe pixel array in the flat panel detector. This is due to the pixelarchitecture in flat panel detectors and unique design of the 2D ASG: 2DASG's septal pitch is in the order of several mms, which is a factor of10 or more than the pitch of photodiode pixel array in the detector. Asalignment of the 2D ASG's septa and detector's pixel array alignment isnot needed, it is feasible to combine a 2D ASG with a flat paneldetector. In contrast to the present invention, 2D antiscattercollimators.

In preliminary experiments with lead based 2D grids that mimickedscatter suppression properties of the proposed design,scatter-to-primary ratio (SPR) was suppressed from 3 to about 0.1 (FIG.11 ). Also it was predicted that signal-to-noise ratio (SNR) improvementprovided by the 2D ASG will be higher than a radiographic ASG withcomparable grid ratio (FIG. 12 ).

Fabrication of a 2D grid from thin enough radio-opaque, or at leastradiation-absorbing, septa is a challenging problem due to limitationsin conventional fabrication and machining methods. However, rapidprogress in additive manufacturing methods for metals made in recentyears now makes manufacturing of 2D ASGs viable [46]. In the currentinvention, advances in additive manufacturing technologies wereexploited (also known as 3D printing) to build one embodiment of the 2DASG. Specifically, Direct Metal Laser Sintering (DMLS) methods wereutilized, which uses a focused, computer guided laser beam to sinterpowdered tungsten to generate the ASG from a 3D computer model. DMLSprocessing of tungsten is a new, but commercially available technology,suitable for both rapid prototyping and serial production of 2D ASGs.

The high mechanical strength of the 2D grid obviates the need forinter-septal spacers. Commonly used in radiographic ASGs to providemechanical support for 1D array of lead septa, inter-septal spacerscontribute significantly to attenuation of the primary x-rays. They alsoimpose a limit on the grid ratio, or grid height, as the thickness ofspacers also increases with grid height, further reducing primarytransmission. Due to absence of inter-septal spacers in the currentinvention 2D ASG, increasing grid height to improve scatter suppressionis not penalized by reduced primary transmission. Another unique aspectof the current invention design is its focused grid geometry; each gridelement will be aligned towards a true point x-ray source in the 3Dcomputer model of the proposed design. Alignment of individual gridelements will be precisely replicated during manufacturing, since theDMLS process is directly driven by the 3D model of the current invention2D ASG. Hence, high primary transmission property of the currentinvention design will be maintained across the active area of the FPD.

As a cumulative effect of the design properties described above, theprimary transmission through the current invention design is predictedto be 20%-50% higher than current radiographic ASGs (FIG. 13 ). Due tothe 2D grid structure, the current invention design with a grid ratio of12 is expected to suppress the scattered radiation intensity by morethan a factor of 2 when compared to a radiographic ASG with the samegrid ratio [1, 2]. Given the high primary transmission and scattersuppression efficiency of the proposed 2D ASG, a robust improvement inlow contrast resolution and CT number accuracy will be achieved in CBCTimages. It is believed that the level of quantitative accuracy providedby the 2D ASG will enable accurate CBCT-based dose calculations in RT.

While two other 2D ASGs have been investigated for mammography, and MDCTsystems in recent years, their applicability to FPD-based CBCT systemsis not feasible. The 2D ASG for mammography systems is fabricated fromcopper [47], which is not radio-opaque, or radiation-absorbing, enoughin the energy range of 60 to 125 keV used in CBCT systems for RT.Moreover, the x-ray lithographic process used in fabrication of themammographic grid is not suitable for fabrication at the physicaldimensions of the current invention, for processing of tungsten. The 2DASG for the third generation MDCT is built from a tungsten polymer blendusing a micro-molding technology [48]. While the ASG for MDCT systemsmay appear similar to the current invention design, parallel (i.e. notfocused) grid elements make it unsuitable for use with a flat detectorCBCT system.

Research Approach

Step 1. Simulate and Build 2D ASG Prototypes Using DMLS Technology, andCharacterize their Signal Transmission Properties.

The goal of Step 1 is to determine the optimal physical properties, orgeometry parameters, of the 2D ASG that would provide the desiredimprovement in CT number accuracy and low contrast resolution within theconstraints of the CBCT system and of DMLS technology. Scatter andprimary transmission properties are primarily determined by fourparameters: grid ratio (i.e. height-to-width ratio of the grid element),septal thickness, grid pitch (i.e. sum of grid element width and septalthickness), and grid height. The impact of these geometry parameters onthe CBCT image quality metrics will be investigated, and optimal valuesfor each geometry parameter will be determined as outlined below.

Simulation of 2D ASG by Monte Carlo calculations: Optimization of ASG'sgeometry covers a large parameter space that might be expensive toexplore by brute-force experimental approach. As outlined in the nextparagraphs, preliminary experiments were performed and a computationalASG model developed to better understand the signal transmissioncharacteristics of 2D ASG. While these studies will be useful to guidethe development of prototypes, a Monte Carlo based ASG model was alsodeveloped to better understand the physics behind 2D ASG's x-raytransmission properties. This way, a better prediction the transmissionproperties of the subsequent prototypes can be made. The EGSnrc package[49] was employed for simulations, which was validated for CBCT [50].Herein, the simulated 2D ASG was validated against the measurementsperformed with the first 2D ASG prototype.

Optimization of grid ratio for desired accuracy in CT numbers: Thescatter component in image signal, commonly quantified by SPR metric,dominates the degradation in CT number accuracy. Siewerdsen andJaffray's work [4] showed that suppressing the SPR down to 0.1 in CBCTprojections kept the overall CT number accuracy within 2-4% of true CTnumbers. An aim was to suppress SPR level to 0.1 with the 2D FASG. Thislevel of accuracy would be sufficient for CBCT based RT dosecalculations (this was evaluated in Step 3).

The level of SPR suppression is directly correlated with the grid ratioof the ASG. In preliminary experiments with a 2D ASG prototype castedfrom lead [1], it was observed that SPR was suppressed from 3 to 0.1with a grid ratio of 12 (see FIG. 11 ). It was predicted that the 2D ASGwith a grid ratio of 12 can suppress SPR even below 0.1 for the sameimaging conditions in [1], due to its higher primary transmissionfraction than the lead based 2D ASG prototype. Thus, the grid ratio of12 was used in the first prototype. Based on the measured SPRsuppression levels and SPR levels predicted by Monte Carlo simulations,a grid ratio range of 8 to 16 will be investigated in the subsequent ASGprototyping cycles to suppress the SPR to 0.1.

Optimization of grid pitch and septal thickness for improvedvisualization of low contrast objects: Percentage of primary x-raystransmitted through the grid plays a key role in the improvement in CNRand SNR metrics (i.e. metrics to quantify the improvement in lowcontrast visualization). Based on the computational 2D ASG model [1],septal thickness and the grid pitch of the first 2D ASG prototype willbe 100 microns and 3 mm, respectively. With this grid geometry, primarytransmission fraction can be improved to >85% (FIG. 13 ), which isexpected to provide significant SNR improvement over radiographic ASG.

Depending on the primary transmission measurements, grid pitches will beexplored down to 5 mm to achieve primary transmission fraction of 85 to90%. Tungsten thickness of 100 microns is more than sufficient to absorbscattered radiation in the diagnostic x-ray energy range. RadiographicASGs typically consist of 25 to 35 microns thick lead septa. FIG. 14Ashows one embodiment of the 2D ASG built for FPD and CBCT applications.FIG. 14B shows the shadow, or footprint, of the 2D ASG appears as adarker grid pattern in the FPD image. X-ray intensity reaching the FPDis lower in the shadow, since 2D ASG's septa partially block primaryx-rays. If this intensity variation is not corrected, it will lead toimage artifacts, and makes the use of 2D ASG unfeasible.

The limit on grid height for physical clearance between the patient andthe CBCT gantry: Given the geometry of linac mounted CBCT systems, themaximum grid height may be limited to 5 cm to maintain sufficientphysical clearance between the patient and the FPD/ASG assembly. Thefirst 2D ASG prototype will have a grid ratio of 12, grid pitch of 3 mm,and grid height of 36 mm. Although the grid height limit also implieslimits on the grid pitch and ratio, it is expected that the optimum gridpitch and ratio will be achieved without increasing the grid heightbeyond the 5 cm limit.

Does the ASG's tungsten footprint introduce dead pixel zones in the FPD?The signal will be reduced in detector pixels underneath the footprintof tungsten septa. If the signal reduction is severe enough, pixelsunderneath the footprint of 2D ASG would appear as dead zones, which maynegatively impact both spatial and low contrast resolution in CBCTimages. Based on computer simulations [1], it is not anticipated thatthis problem will occur due to the relatively thin septa with respect todetector pixel size; less than 5% of all detector pixels are expected tohave signal reduction by more than 50% due to proposed FASG's footprint.Even if these pixels are treated as “dead” pixels, the signal at deadpixel locations can be easily recovered by the remaining 95% of pixelsusing interpolation methods.

Construction of 2D ASG prototypes using DMLS technology: The ASGprototypes used in Step 1 were initially 10×10 cm² in size to minimizeproduction costs. This size will be more than sufficient to measure theprimary and scatter transmission characteristics of the design.

Experiment setup: Throughout this invention, ASGs were evaluated in aclinical, linac mounted CBCT system (TrueBeam, Varian Medical Systems,Palo Alto, Calif.). This approach enabled evaluation of the 2D ASG usingclinical imaging geometry, system components, and protocols [51]. Twoutilized TrueBeams have “research mode” capability, which allowed fullcontrol the CBCT system parameters, and image processing outside theclinical CBCT system. In all experiments through Step 1 to 3, directlymounted the prototype 2FASGs on the FPD, and was employed with the“offset detector” CBCT geometry (also known as half-scan geometry) inTrueBeam linacs. It is envisioned that half-scan geometry is the optimalchoice for the 2D ASG as it provides the largest field of view. WhileVarian's CBCT platform was selected for this project, the currentinvention design can be adapted to other CBCT systems, such as XVI(Elekta Oncology Systems, Norcross, Ga.).

Experimental details: With each ASG prototype, the scatter transmissionfraction, average primary transmission fraction, and SPR suppressionlevels were characterized. Measurement of primary and scattertransmission will be performed in CBCT projections using establishedtechniques, such as the beam-stop method [52]. The full detector area(30×40 cm²) will be exposed to simulate imaging conditions in clinicalsystems, and SPR suppression characteristics of each prototype will beinvestigated for patient thicknesses up to 50 cm, where the goal is tolimit SPR to 0.1 even in high SPR environments. The signal reduction dueto ASG's footprint will also be evaluated and results will be comparedto the Monte Carlo model. Based on the scatter suppression and primarytransmission properties of the first prototype, the geometry of thesubsequent prototype will be generated to minimize the discrepancybetween the desired and measured ASG performance metrics. It isestimated that 4 to 6 cycles of ASG prototype construction andcharacterization will be performed to determine the optimal ASG geometryparameters.

Step 2. Evaluate the Impact of 2D ASG on CBCT Image Quality.

In Step 2, improvements in CBCT image quality provided by the 2D ASGwere characterized. The prototype to be used in Step 2 will be based onthe optimal ASG geometry parameters determined in Step 1, and it willcover the entire active area of the FPD to reflect the image qualityimprovements in clinical CBCT systems.

Experiment setup: As in Step 1, experiments will be performed in “offsetdetector” CBCT geometry using the CBCT system in TrueBeam linac. Toemulate clinical CBCT imaging protocols, the bowtie filter was employedto demonstrate the combined effects of bowtie filter and 2D ASG onscatter suppression and CBCT image quality. CBCT images will bereconstructed using the FDK filtered back-projection algorithm [53].Beam hardening [54], image lag [55], and ring artifact correction [56]algorithms will be implemented in the post-processing stage of CBCTprojections to further improve CT number accuracy.

Is 2D ASG's weight an issue for the CBCT gantry? For the geometryparameter ranges outlined in Step 1, the maximum weight of detector andgrid assembly will be 18.1 kg, which is below the CBCT detector arm'spayload limit of 20 kg in the TrueBeam linac [51].

Correction of 2D ASG's footprint in CBCT projections in the presence ofgantry flex: Linac gantry and the CBCT imaging arms flex (or sag)slightly due to gravity as they rotate around the patient. Projectedlocation of x-ray source on the FPD will vary due to flex, and footprint(x-ray shadow) of the ASG's septa projects to slightly differentlocations on the FPD as a function of gantry angle.

The characteristics of linac-CBCT gantry flex are reproducible, or“systematic”, as a function of gantry angle, which makes the correctionof 2D ASG's footprint feasible [57-61]. A gantry angle-specific flatfield correction method was first implemented, which has been shown tosuccessfully correct the footprint of radiographic ASGs in the presenceof gantry flex [40, 62]. This process correction can easily be automatedand needed calibration images can be acquired during routine monthlylinac-CBCT QA sessions. Although gantry flex is quite reproducible overextended periods of time, small “random” variations in gantry flex exist[57], which may lead to ring artifacts in CBCT images. To account forthis, a post-reconstruction ring artifact suppression algorithm, [63].Such algorithms have already been successfully implemented in clinicalCBCT systems will be implemented [51, 64]. It is worth mentioning thatgantry flex will be insignificant in a ring based linac gantry (ratherthan c-arm like gantry), which may negate the need for the ASG footprintcorrection methods described above. Ring based linacs have been readilyoffered by a linac Vendor (Vero, BrainLab, Munich, Germany).

Studies to be performed: The studies in Step 2 will be an extensive setof CBCT image quality characterization experiments. The magnitude andeffects of scattered radiation on image quality strongly depend on thesize, shape, and composition of the imaged anatomy. Thus, we will employa comprehensive set of anatomy and patient size specific phantoms toevaluate the imaging performance. To assess the effect of anatomy, head,thorax, and pelvis anthropomorphic phantoms will be utilized. To assessthe effect of phantom size and measure various image quality metrics,the Catphan phantoms will be utilized (The Phantom Laboratory, Salem,N.Y.). Catphan can be fitted into various size “body” phantoms mimickingpelvis/abdomen anatomy. With body phantoms, we will simulate patientsizes up to 55 cm lateral and 45 cm anteroposterior separation. Catphanphantoms were evaluated at 3 different body sizes. Imaging technique(tube kVp, mAs) will be determined based on the clinical protocols inTrueBeam.

For each phantom, the following 3 sets of images: 1) CBCT withradiographic ASG, 2) CBCT with the proposed 2D ASG, 3) MDCT using the CTsimulator was acquired. MDCT image represents the gold standard in imagequality, thus, image quality of CBCT images will be benchmarked againstthe MDCT images in each imaging experiment. The CT dose index (CTDI) wassimilar for both CBCT and MDCT to suppress the impact of imaging dose onimage quality.

Established methods and metrics were employed for image qualityevaluations [4, 28]; contrast to noise ratio (CNR) was measured, asurrogate for improvement in soft tissue visualization, using contrastobjects embedded in phantoms. Statistics of CT numbers (mean andstandard deviation) were measured in 10-15 different regions of interestto assess the accuracy of CT numbers. In experiments with Catphanphantoms, modulation transfer function (MTF), were also measured inaddition to the metrics mentioned above.

Measurements in CBCT and MDCT images will enable us to preciselycharacterize the difference in image quality metrics across 3 differentsystems (i.e. radiographic ASG, proposed 2D ASG, and MDCT) for differentanatomies and patient sizes. To identify the similarities (anddifferences) among different systems (2D ASG vs. MDCT and 2D ASG vs.radiographic ASG), the differences in image quality metrics will testedfor statistical significance.

Step 3. Validate Improvements in the Accuracy of Radiation Therapy DoseCalculations.

Achieving clinically acceptable dosimetric accuracy in CBCT-basedtreatment plans is an important milestone for the invention. In Step 3,whether improved CBCT image quality is sufficient for treatment dosecalculations in megavoltage external beam RT will be tested.

Experiment setup: Step 3 will employ the image sets acquired in Step 2for treatment plan dose calculations. The only new imaging experimentsin Step 3 will be with CT number to electron density (CT-to-ED)phantoms; to be able to calculate dose in CBCT or CT image, CT numbersare converted to either electron or mass density values in the treatmentplanning system. To achieve this, CBCT and MDCT images of CT-to-EDphantoms will be acquired, and CT-to-ED tables will be established [65].

Evaluation of dosimetric accuracy in CBCT: In the Radiation OncologyClinic, dosimetric accuracy analysis has been routinely performed forboth clinical cases and research projects [66-68]. The same tools andmethodology will be utilized, such as point-by-point dosimetriccomparisons and analysis of dose volume histograms (DVH) [69, 70], toquantify the dosimetric accuracy provided by 2D ASG. As in Step 2,MDCT-based treatment plans and dose distributions are considered thegold standard. For a given phantom, CBCT-based dose distributions willbe directly compared to the MDCT-based dose in each treatment plan.Below, the general method for dose accuracy evaluations for eachanatomical site or phantom is provided.)

For each phantom, all 3 sets of images acquired were utilized in Step 2(i.e. CBCT with radiographic ASG, CBCT with 2D ASG, MDCT). To generateradiation treatment plans, images will be imported into a clinicaltreatment planning system (Eclipse, Varian Medical Systems, Palo Alto,Calif.). 2) Depending on the anatomy depicted in the phantom, relevanttargets and organs at risk structures were generated. 3) In each CBCTimage, treatment plans will be generated using both 3D conformal andintensity modulated radiation therapy techniques. Attention will be paidto emulate realistic treatment plan scenarios. A clinically employedlist of organs at risk and target dose constraints was used to optimizethe plans. Once the plan was optimized, final treatment plan and dosecalculation parameters were copied to the corresponding MDCT image set,and the dose will be recalculated to obtain the “ground truth” dosedistributions. This way, the dose difference between CBCT and MDCT-basedplans will only be due to differences in CT numbers of image sets. 4) Asin clinical plans, DVHs will be calculated for organs at risk andtargets in both CBCT and MDCT based plans. A set of site-specific DVHmetrics were selected to assess the dose distributions as in clinicalcases (such as dose covering 95% of the target volume, mean lung doseetc.). Also, point-by-point dose comparisons will be performed betweenthe CBCT and MDCT dose distributions. The statistical significance ofdifferences in point-by-point dose distributions of CBCT and MDCT-basedplans was tested. Statistical analysis will allow us to identifyanatomies and imaging conditions that lead to dosimetric differencesbetween CBCT and MDCT 5) From a clinical perspective, evaluation of“clinical” significance of dose differences by expert clinicians is amore relevant approach (e.g. Statistically significant differences atvery low doses that occur close to skin may not be considered clinicallysignificant). The differences in DVH metrics and dose distributions ofCBCT and MDCT-based plans were evaluated from a clinical standpoint.With this approach, we Step to map out the anatomical sites, patientsizes, and treatment techniques that provide clinically acceptabletreatment plans with the use of the proposed 2D ASG.

The impact of improved CBCT image quality may be moderate in the contextof current clinical use of CBCT, which has been limited to in-roompatient setup corrections. However, it is believed that the impact ofimproved CBCT image quality is best viewed in the context of enablingnew paradigms in clinical practice, such as adaptive radiation therapy(ART). In-room CBCT imaging plays a key role in various components ofART framework, such as monitoring of normal tissue and tumor changes,and assessing dosimetric consequences during the course of treatment.Although much work has been done on the key tools for ART (e.g.deformable image registration, dose accumulation, and tissuesegmentation tools), such tools don't work reliably with today's CBCTimages due to lack of soft tissue visualization and inaccuracy of CTnumbers. Furthermore, assessment of tumor response by a physician, anddecision for treatment modification are also challenged by poor CBCTquality.

The lack of CBCT utilization in ART is a major roadblock to widerclinical implementation of ART, and poor CBCT image quality due toscatter is a major contributor to the problem. When successfullyimplemented, the current invention 2D ASG design represents abreakthrough in CBCT image quality improvement, in turn enablingtreatment response monitoring and adaptation methods by usingincreasingly common linac-mounted CBCT systems. Also, CBCT is utilizedin clinical applications outside radiation oncology, such astransarterial chemoembolization and maxillofacial surgery, where poorimage quality due to scatter is considered a major drawback. Successfulcompletion of this project will impact such applications of CBCT outsideradiation therapy.

The current invention proposes both a novel architecture and use of anovel 3D printing technique, known as Direct Metal Laser Sintering(DMLS), for tungsten. The current invention grid architecture representsa fundamental departure from existing 1D antiscatter grid designs. Allcurrent antiscatter grids in CBCT are based on the 1D Potter-Bucky griddesign, which is unsuitable for design of 2D antiscatter grids. Thecurrent invention architecture reduces the grid pitch by more than anorder of magnitude, and increases the septal thickness by a factor of 3to 5. Importantly, tungsten has been chosen to provide mechanicalstrength, higher radio-opacity, and eliminate inter-septal spacers. Itis the use of DMLS which permits realization of a 2D grid and allowsfabrication with tungsten. These deliberate design choices have beenmade to achieve both favorable imaging performance and fabricationfeasibility.

Septal Shadow Aft Fluence Modulation

While Cone Beam Computed Tomography (CBCT) is routinely used as a 3Din-room imaging modality in radiation therapy, its poor soft tissuevisualization and lack of material composition information is believedto prevent implementation of improvements in radiation therapy, such asCBCT-based treatment monitoring and treatment plan adaptations.

Currently used x-ray imagers utilize energy integrating detectors thatcreate a 2D map of total energy deposited within an x-ray sensor.Consequently, information about the number of x-rays and the energy ofeach x-ray absorbed in the sensor is lost. This shortcoming of energyintegrating detectors deteriorate image quality, prevent implementationof new technology for x-ray imaging, and increase radiation doses topatients during imaging procedures.

To address these shortcomings of energy integrating detectors, photoncounting detectors were developed. For example, photon countingdetectors can quantitate the energy of individual x-rays and count eachx-ray that interacts with a detector. Although it is not necessary tounderstand the mechanism of an invention, it is believed that photoncounting detectors will be integrated into future computed tomographyand radiography systems, especially those contemplated herein.

However, currently available photon counting detectors have severaltechnical challenges that interfere with their utilization in clinicalimaging systems. For example, in clinical imaging scenarios, x-rayincident flux on a detector can be quite high (e.g., up to 10⁸-10⁹x-rays per mm² per second). At this high x-ray flux photon countingdetectors are unable to accurately count the individual x-rays (e.g.,commonly referred to as a “pulse pile-up”) and, in addition, thedetected number of x-rays can be significantly different than the actualnumber of x-rays (e.g., a significant error rate). FIG. 23 . Although itis not necessary to understand the mechanism of an invention, it isbelieved that this problem prevents the reconstruction of high qualityx-ray images.

In one embodiment, a 2D anti-scatter grid comprises series of holesseparated by a radio-opaque, or at least radiation-absorbing, septa(dashed red line). FIG. 24A. In one embodiment, the holes comprise anyshape, but are preferably either square or hexagonal. See, FIG. 1 . Animage profile of one of these hexagonal hole series demonstrates thevariations in x-ray flux depending on the pixel position within thehexagonal hole. FIG. 24B (orange line). It can be seen that a darkerregion of the hexagonal hole series in FIG. 24A corresponds to a 2D gridfootprint (e.g., septal shadow) indicating a lower x-ray flux orintensity. Consequently, the pixels within the septal shadow encounter asignificantly reduced x-ray flux as compared to pixels located withinthe 2D grid's radio-opaque, or at least radiation-absorbing, walls.(e.g., lighter regions of FIG. 24A). In one embodiment, theradio-opaque, or at least radiation-absorbing, walls include a materialincluding, but is not limited to, tantalum, tungsten or lead. In otherwords, a 2D grid acts as spatial modulator of x-ray flux, havingspecific regions of high or low intensity.

Dual energy CT techniques, such as material decomposition and virtualmonoenergetic imaging, can play a role in mitigating such shortcomingsof CBCT imaging. However, dual energy imaging in flat panel detector(FPD) based CBCT systems faces major challenges: since image acquisitionrates of FPDs is a factor of 30-60 slower than conventional CTdetectors, consistency of CT numbers needed for dual energy processingcannot be achieved in the presence of organ motion. One current dualenergy method, projection domain, has proved challenging to implementwith FPDs. This is because high and low energy projections cannot beacquired at similar source-detector positions. Additionally, asufficient number of dual energy projections may not be acquired due toslow image acquisition, which may lead to under-sampling problems. Toaddress these problems, one embodiment of the present inventioncontemplates a photon counting and energy resolving detector that isneeded in CBCT imaging. In contrast to FPDs, photon counting detectorsacquire high/low energy projections simultaneously and therebyeliminates organ motion in such projection pairs. Hence, dual energyCBCT can be enabled via implementation of projection domain dual energyprocessing methods.

In one embodiment, the present invention contemplates a photon countingdetector that improves low contrast resolution. Although it is notnecessary to understand the mechanism of an invention, it is believedthat low contrast resolution improvement is due to an immunity toelectronic noise and/or higher x-ray detection efficiency than FPDs. Afurther advantage is that photon counting detectors negate a need forrapid x-ray tube voltage-switching, where adequate dose allocation andspectral separation between in high/low energy images can be challengingto achieve.

The data presented herein demonstrate that the contemplated 2Dantiscatter grid and aft fluence modulation method have overcome twomajor roadblocks known in the art. First, a “scattered radiation”intensity can exceed a primary intensity by a factor of 5 to 8, whichseverely biases the counts in high/low energy bins and makes dual energyimaging unfeasible. Second, a “limited count rate” (e.g., a “pulsepile-up” effect) is a major limitation of photon counting detectors,where the quantity and energy of x-rays cannot be accurately resolved athigh x-ray fluence imaging conditions.

In one embodiment, the present invention contemplates devices andmethods that overcome the above discussed problem of “scatteredradiation” when using photon counting detectors. As detailed above, thepresent invention contemplates a two-dimensional antiscatter gridcomprising a grid architecture that fundamentally differs with respectto radiographic 1D antiscatter grids. Although it is not necessary tounderstand the mechanism of an invention, it is believed that a 2D ASGas contemplated herein provides superior scatter suppressioncharacteristics when compared to conventional radiographic grids andconventional scatter correction methods.

In one embodiment, the present invention contemplates devices andmethods that overcome the above discussed problem of “limited countrate” when using photon counting detections. In one embodiment, thepresent invention contemplates a method comprising an aft fluencemodulation. Although it is not necessary to understand the mechanism ofan invention, it is believed that an aft fluence modulation utilizes aseptal shadow of the 2D ASG where detector pixels underneath a grid'ssepta are exposed to significantly lower x-ray fluence than adjacentpixels in the center of 2D grid's cells.

Although it is not necessary to understand the mechanism of aninvention, it is believed that in one embodiment, a 2D ASG grid isconfigured as an x-ray flux modulator. In one embodiment, a plurality ofdetector pixels are configured underneath a septal shadow such that theincident flux is reduced and x-rays can be accurately counted at eachspecific pixel. See, FIG. 25 (yellow arrows). Furthermore, X-ray countsin other pixels (i.e. pixels in the center of the holes) can becorrected by using the accurate count information in pixels underneaththe grid's footprint or septal shadow.

Hence, an improved count accuracy and energy information may beextracted from pixels within the septal shadow. In one embodiment, theimproved count accuracy and energy information is processed to correcthigh fluence region pixels that are biased by the “pulse pile-up”effect. Besides improving count rate capability, it is believed thatseptal shadows of a 2D ASG as contemplated herein can also reduce chargesharing among neighboring pixels, a known source of energy resolutiondegradation in photon counting detectors.

While various fluence modulation methods were previously proposed, allwere based on pre-patient attenuators, such as dynamic bow tie filters.However, dynamic filters require a predictive algorithm and priorpatient information to determine a patient specific modulation pattern,and can be electromechanically complex to achieve the desired fluencemodulation. The presently disclosed configuration of pixels within theseptal shadow of a 2D ASG as contemplated herein support an aft fluencemodulation method that is not predictable from known dynamic pre-patientattenuators as a stationary 2D ASG is mechanically simple and priorpatient knowledge is not required.

In one embodiment, the present invention contemplates a 2D ASGcomprising a photon-counting detector wherein a plurality of pixels arewithin a septal shadow of the 2D ASG. In one embodiment, the septalshadow reduces x-ray scatter intensity. In one embodiment, the septalshadow reduces spatial variation in x-ray fluence. In one embodiment,the pixels within the septal shadow improve energy resolutionperformance. The data provided herein experimentally validated aftfluence modulation and characterized an improvement in count rate andenergy resolution performance. Also evaluated are fluence conditionsobserved in clinical CBCT imaging.

The data provided herein will experimentally evaluate scatter reductioneffects provided by 2D ASG. These data will be fabricated on 2D ASGprototypes and integrated with a photon counting detector.

Photon Counting Detectors

In one embodiment, the present invention contemplates an x-ray devicecomprising a 2D antiscatter grid and a hybrid flat panel detector,wherein the detector comprises a plurality of photon counting detectors.In one embodiment, the photon counting detector comprises a substrate(e.g., a cadmium telluride substrate) positioned between a cathode plateand an anode plate. In one embodiment, the anode plate comprises aplurality of pixels. Although it is not necessary to understand themechanism of an invention a photon (e.g., an x-ray photon) impingingupon the substrate is detected by a pixel, such that it emits signalproportional to the energy of each photon (blue arrow). See, FIG. 26 .In one embodiment, the photon counting detectors are positioned below aradio-opaque, or at least radiation-absorbing, area of the 2D ASG.

Current CBCT systems based on flat panel detectors (FPD) exhibit a poorlow contrast visualization and lack CT number accuracy, whichconstitutes a barrier to implementation of treatment strategies inradiation therapy. When compared to MDCT detectors, amorphous siliconFPDs have significantly higher electronic noise, lower quantumefficiency, and lower digitization range [1]. These problemscumulatively contribute to degradation of low contrast objectvisualization in CBCT images [2]. One drawback of FPD is a relativelylow projection acquisition rate (15-40 projections/sec) as compared toMDCT (2000-3000 projections per sec) [3]. Lower frame rates, combinedwith the safety-mandated slow gantry rotation speeds of linac gantries(60 secs per rotation) result in the presence of organ motion in CBCTimages, which in turn prevents implementation of dual-energy CT (DECT)techniques. Among clinically available DECT acquisition techniques, onlyrapid kVp switching would be potentially applicable to FPD based CBCT[4, 5]. In this method, the tube voltage is switched between high andlow kVp values (e.g. 80 and 140 kVp) to acquire high and low energyprojections at a periodicity corresponding to the frame rate of thedetector. However, due to the limitations of FPDs described above,low/high energy projections cannot be acquired fast enough, or atsufficiently similar gantry angles. As a result, dual energy processingmay not be directly performed on low/high energy projections pairs,which has been reported to require pairs to be acquired at the samegantry, or source/detector position [4].

A potential alternative to DECT acquisition techniques is dual energyprocessing in the image domain (i.e. after the reconstruction of highand low energy CBCT images). However, in the clinical context of CBCT,this is complicated by organ motion. In addition to severely reducing CTnumber accuracy, organ motion may introduce pronounced motion artifactsin high and low energy reconstructions, which prevents accurateestimates of monoenergetic or material-specific CBCT images.

In one embodiment, the present invention contemplates a hybrid flatpanel detector comprising a plurality of photon counting detectors.Although it is not necessary to understand the mechanism of aninvention, it is believed that photon counting detectors de-coupledual-energy processing from organ motion artifact. For example, it isbelieved that each photon counting detector projection contains bothhigh and low energy images acquired at the same time [6]. Hence, eachprojection is free of organ motion, and dual energy processing can beperformed directly on each projection individually. Subsequently,material specific or monoenergetic CBCT images can be reconstructed.While effects of organ motion is present in a material-specific CBCTimage (due to the slow rotation of CBCT gantry) it will not interferewith projection domain dual energy processing.

Photon counting detector based CBCT systems face two major barriers:

-   -   a) High scattered radiation intensity is an inherent problem of        CBCT [2, 9]. Scatter contaminates the energy spectrum such that        dual energy processing may not be performed accurately. As such,        it is currently believed that DE CBCT research is limited to        imaging of small objects, such as extremities or small animals,        where scatter intensity is relatively lower [5, 10, 11]; and    -   b) CBCT count rate limitation due to pulse pile-up. Pulse        pile-up occurs when two or more x-ray events are absorbed in the        x-ray sensor with a very small temporal separation, and hence        register as a single x-ray event leading to x-ray counts being        underestimated [6, 12]. The incident x-ray fluence in CBCT can        be as high as 10⁷-10⁸ x-rays per second/mm². Such incident count        rates are still high enough to cause considerable pile-up and        count rate losses.

In photon counting MDCT, pulse pile up is particularly a big problem asmaximum fluence incident on the detector can be in the order of 108-109counts per second/mm2 and, as a result, x-ray counts are under-measureddue to pulse pile-up. See, FIG. 27 . In CBCT, the incident x-ray fluenceis much lower than MDCT, about 107 counts per second/mm2. It is believedthat this low fluence may be due to a factor of 30 to 100 lower tubecurrents and a consequent a slower gantry rotation time (e.g., 60seconds in CBCT vs 0.5 seconds in MDCT). However, such incident countrates are still high enough to cause considerable pile-up and count ratelosses.

In one embodiment, the present invention contemplates both devices andmethods to solve the CBCT problems regarding both the scatteredradiation and limited count rate problems for photon counting detectors.In one embodiment, these devices and methods comprise photon countingdetectors. In one embodiment, these devices and method comprise dualenergy techniques. In recent years, many clinical studies have beenpublished on improving soft tissue visualization by synthesizing low keVmonoenergetic images, and the benefits of material decomposition, allenabled by DECT [13-16]. It is currently believed that translation ofDECT techniques to a CBCT domain requires overcoming the above definedCBCT image quality barriers. In one embodiment, the present inventioncontemplates a photon counting method based upon dual energy processing.

In some embodiments, a CBCT photon counting approach has advantages overenergy integrating CBCT; i) as x-rays are counted, photon countingeliminates a readout noise problem associated with FPDs; ii) an X-raycounting approach provides much higher dynamic range than energyintegrating FPDs; iii) X-rays are directly converted to electricalcharge, i.e. scintillators are eliminated, and thus, sensor thicknesscan be increased to 2 mm or more to improve quantum efficiency above90%, without sacrificing spatial resolution [6].

In one embodiment, the present invention contemplates a 2D ASG as atwo-dimensional array of through-holes separated by radio-opaque, or atleast radiation-absorbing, septa. See, FIG. 1 . In one embodiment, eachthrough-hole is aligned, or focused, towards an x-ray source. Moreover,the 2D ASG contemplated herein lacks an interseptal spacer as anoptimized grid design that provides a high primary transmission. In oneembodiment, the present invention contemplates a 2D ASG comprising aCBCT flat panel detector. FIG. 28A and FIG. 28B. Although it is notnecessary to understand the mechanism of an invention, it is believedthat a 2D ASG reduces scatter contamination to levels that has not beenachievable with any existing CBCT scatter mitigation approaches. It isfurther believed that a two-dimensional array of septa as presentedherein is a more efficient configuration to stop radiation scatter thanany one-dimensional septa array currently used in CBCT radiographicimaging ASGs.

It has been reported that a 2D ASG with a grid ratio of 8 provided afactor of 3 to 6 lower SPR than a radiographic 1D ASG with a grid ratioof 10. FIGS. 29 , and [17]. This is further evidenced by an improvedquality of CBCT images when using a 2D ASG with a hybrid flat panelphoton counting detector versus a 1D ASG with radiation gatheringdetectors and a radiation gathering detector without an ASG. FIGS. 30A,30B and 30C.

A reduced scatter fluence by a 2D ASG is also expected to reduce thephoton “pulse pile-up” problem. A scattered radiation to primaryradiation ratio measured using a 30 cm phantom thickness exceeds four(4), which indicates that scattered radiation increases x-ray intensityincident on the detector by a factor of 4. Since scattered x-raysincrease the x-ray fluence incident on a detector, this phenomenoncontributes to a pulse-pile up problem. FIG. 29 . However, with a 2D ASGas contemplated herein, a scattered radiation-to-primary radiation ratiois reduced to 0.31. Consequently, a 2D ASG configuration leads to analmost factor of four reduction in scattered radiation intensityincident on a detector, and hence, reduces photon pulse pile-up. Withradiographic 1D ASGs, this effect will be less pronounced as theytransmit high fraction of scattered x-rays. Software based scattercorrection strategies cannot reduce the scatter induced pulse pile-up,as scatter correction is performed after the detection of scatteredx-rays by the x-ray sensor.

With respect to a 1D ASG, scatter suppression efficiency of 2D ASG wasparticularly better at higher scatter intensity environments; as theangular distribution of radiation scatter is larger for thicker objects,septa configured in two dimensions is more efficient in stoppingscattered x-rays than one-dimensional a septa configuration.

As mentioned above, a 2D ASG provides a higher primary radiationtransmission than 1D ASGs. It is believed that the higher primaryradiation transmission is due to: i) an absence of interseptal spacers;and ii) a larger grid pitch (2.9 mm). In one embodiment, a 2D ASG ascontemplated herein comprises an 85% primary transmission (averaged overthe entire area of 2D ASG), whereas a conventional 1D ASG with fiberspacers has a 71% primary transmission.

The implementation of 2D ASGs in the context of photon countingdetectors and dual energy imaging is believed to have advantages over 1DASGs and 2D ASGs configures with energy gathering detectors. Theseadvantages include, but are not limited to: a) a reduction of energyspectrum contamination; b) a reduction in scattered radiation fluencecontamination that is characteristic of x-rays emitted from tungstengrid septa; c) a reduction in radiation backscatter from photon-countingdetector on an incident x-ray energy spectrum; and d) preferred gridsepta positioning with respect to photon counting detector pixelpositions to optimally process x-ray fluence and energy spectrum.

In one embodiment, the present invention contemplates an aft fluencemodulation method comprising a 2D ASG and a photon counting detector forimproving radiation count rate and energy resolution. In one embodiment,a 2D ASG footprint creates a plurality of septal shadows. In oneembodiment, a 2D ASG comprises a plurality of photon counting detectorscomprising pixels are located underneath 2D ASG vertical septa. As such,these photon counting detector pixels are exposed to a significantlylower fluence than pixels in the center of the 2D ASG through-holes.See, FIG. 24 and FIG. 25 . Thus, pixels in septal shadows are lesslikely to be affected by the pulse-pile up problem. In high fluenceregions of a projection, such as skin-air boundaries, pixels in thecenter of through-holes exhibit photon pulse pile-up, and they will beexcluded from the projection image. Image signal in these locations maybe estimated by image signals from septal shadows surrounding thatlocation. As the grid pitch will be in the order of 1-3 mm, we believethat interpolation may be a feasible approach to estimate image signalin the center of the through-holes by using image signal in surroundingseptal shadows.

As the above-described method will be used only in high fluence regionsthat exhibit photon pulse pile-up. However, to solve this problem alsoinvolves locating high fluence regions within a projection. In oneembodiment, an aft fluence method comprises detecting regions exhibitingphoton pulse-pile up. FIGS. 31A, 31B, and 31C. For example, SPR ratio ofcounts in pixel 5 (high fluence region) as compared to pixel 3 (lowfluence region) is not constant, i.e. will change as a function ofincident fluence. The count ratio of pixel 5 to pixel 3 at low incidentfluences is expected to be higher than the ratio at high fluences. Thisis due to the fact that pixel 5 will exhibit more count loss, or have arelatively lower counts, due to pulse pile-up. The count ratio betweenpixels 5 and 3 can be established for varying fluence levels, and thiscalibration data can be utilized as a “pulse pile-up detector” duringCBCT scans.

In one embodiment, the present invention contemplates an aft fluencegrid configuration. For example, as both grid geometry and pixeldimensions play a role in fluence modulation in septal shadows, variousgrid septa configurations are contemplated. FIGS. 31A and 31B. In oneembodiment, septal thickness is increased to further reduce fluencewithin septal shadows. This way, a larger portion of each pixel isshadowed by thicker septa, where a small portion pixel surface isexposed to incident x-rays. However, unobstructed areas of pixels arelocated on or near the edges of pixels, and a charge cloud created byx-ray absorption in this region is more likely to spread ontoneighboring pixels, which may lead to spectral degradation due to chargesharing. A more optimal solution to reduce fluence in septal shadows isto employ grid septa with a footing. FIG. 31B. In one embodiment afooting has a height of approximately between 0.3 to 2 mm. Although itis not necessary to understand the mechanism of an invention and it isbelieved that the footer attenuates the radiation beam, rather thanfully obstructing x-rays. By using a septal footing, fluence reductioncan be better controlled, and a large portion of the pixel surface maybe exposed to incident x-rays, potentially reducing the charge sharingproblem. With this approach, fluence within septal shadows can bereduced by an order of magnitude lower when compared to fluence at thecenter of septal through-holes.

Additive manufacturing methods are used to create grid geometries andsepta configurations. Powder bed laser melting (PBLM) method is used tofabricate our 2D ASG prototypes. PBLM for tungsten is compatible with anadditive manufacturing process, as challenges due to the high meltingpoint of tungsten have been addressed in recent years. In PBLM, tungstenpowder is spread over a built platform, and a high power laser beamtraces and melts the tungsten powder based on the CAD design. The gridis built layer by layer, by lowering the built platform, adding a newlayer of tungsten powder, and repeating the laser tracing process.Finished products are accurate in dimensions within 20 microns, andminimum septal thickness of 100 microns has been achieved. PBLM provideslarge design freedom, which enables fabrication of complex gridgeometries, such as focusing of individual through-holes, spatiallyvarying grid pitch and height. PBLM may also enable fabrication of ourproposed septa with footing for Aft Fluence Modulation. PBLM is ascalable manufacturing process, which is suitable for both rapidprototyping and serial production of 2D ASGs.

To reduce high x-ray fluence, dynamic beam attenuators were proposed andinvestigated by various research groups in the context of MDCT [18, 19].Such attenuators are placed between the x-ray tube and the patient,where fluence is modulated spatially and temporally, to reduce fluencein less attenuating regions of the patient. Some examples of thisapproach are dynamic bow tie filters and beam-attenuating rods that movein and out of x-ray beam.

Implementation of such techniques may have major limitations,particularly in the context of CBCT. First, dynamic beam filters areaimed for fan beam CT geometry, where fluence modulation is needed inone dimension. In contrast to MDCT, fluence modulation in CBCT is neededin two dimensions, and the desired fluence pattern may not be achievedby a linear array of beam attenuators. For example, in thorax, highfluence regions due to lungs are surrounded by low fluence regions dueto mediastinum and chest wall. Such high fluence “islands” arechallenging to compensate using a linear array of dynamic beamattenuators. Second, determination of modulation pattern requires priorknowledge about the patient attenuation characteristics. If a patient'sprior CT scan is employed for this purpose, differences in patientposition and geometry can lead to regions with under-, over-modulation.Third, since dynamic beam filters are placed close to the x-ray tube,any filter positioning errors will be magnified in the detector plane.The effects of gantry sag and associated effects on positioning errorsare magnified as well. The mismatch between the expected and actualposition of attenuators will lead to high gradient spikes and dips influence, which may cause severe artifacts in images. Lastly, proposeddynamic attenuators are electromechanically complex. Fabrication of suchdevices remains to be an area to be investigated.

In the presently contemplated Aft Fluence Modulation method, theelectromechanical complexity of dynamic attenuators is not present, asfluence is always modulated, in two dimensions, by the 2D grid patternattached to the detector. The methods described herein do not requireprior information about patient attenuation characteristics either.Since the 2D ASG is placed close the detector plane, any adverse effectsof gantry sag or flex are much less pronounced with respect topre-patient beam attenuators.

EXAMPLES

The following examples are provided in order to demonstrate and furtherillustrate certain preferred embodiments and aspects of the presentinvention and are not to be construed as limiting the scope thereof.

Example 1

In one embodiment, the invention relates to a two-dimensionalantiscatter grid, wherein the two-dimensional antiscatter grid isinstalled on top of a flat panel detector protective cover. In oneembodiment, the two-dimensional antiscatter grid is described in FIG. 20, FIG. 21 , and FIG. 22 . In one embodiment, some basic x-raytransmission characteristics of the two-dimensional antiscatter grid aredescribed in Example 2. In one embodiment, the hybrid flat paneldetector is described in Example 3 & Example 4. In one embodiment, thetwo-dimensional antiscatter grid unique qualities and grid parametersare described in Example 5 & Example 6. In one embodiment, a method tocorrect scatter using a 2D antiscatter grid is described in Example 7.In one embodiment, an illustration of a pixelated scintillator is foundin FIG. 20 , FIG. 21 , and FIG. 22 . A more expansive description of oneembodiment of the current invention is described in Example 2. Previousexplorations of antiscatter grid prototypes are described by Altunbas[1, 71].

A. A Flat Panel Detector that is “Directly” Integrated with 2D ASG,Referred as “Hybrid” FPD

While 2D ASG's purpose is to block scattered (or contaminant) x-rays, italso partially blocks primary x-rays (i.e. desired x-rays that form thex-ray image). 2D ASG's septa create a shadow that appears as a gridpattern in FPD images. This shadow is due to the lower x-ray intensityreaching the FPD, as the 2D ASG's septa partially blocks the primaryx-rays. 2D ASG's shadow, i.e. variation in primary x-ray intensity,creates two major problems from image quality point of view: 1) The 2DASG's shadow will cause severe artifacts in images, which makes the useof 2D ASG unfeasible. Therefore, the shadow of the 2D ASG must becorrected by means of software or hardware methods. 2) Even if the2DASG's shadow is corrected, 2D ASG continue to block primary x-rays,and reduced intensity of primary x-rays will reduce image quality: Lowerthe intensity of primary x-rays reaching the FPD, the noisier the imagegets. This problem can be compensated by increasing the x-ray dose tothe patient; this is not, however, a preferred solution. 3) When 2DASG/FPD combo is employed in a CBCT system, the position of 2D ASG'sshadow will shift slightly in relation to the pixels in the FPD, as theCBCT gantry rotates around the patient during image acquisition. As aresult, grid shadow correction methods may not work at all, or maycorrect grid shadows sub optimally.

The shift in grid shadow position is due to the CBCT gantry flex, whichis caused by forces exerted by gravity on the gantry during gantryrotation. As a result, the x-ray source will change position withrespect to 2DASG/FPD combo, and the 2D ASG's shadow will project to aslightly different location. If the 2D ASG's shadow shifts during imageacquisition, correction of the shadow may become quite challenging.

Possible Solutions:

A. 2D ASG's shadow should be minimized as much as possible. To achievethis, it is possible to directly place the 2D ASG on thescintillator/pixel array in the FPD. The shadow size is a function ofthe “distance” between the top surface of the 2D ASG and the FPD pixelarray. Smaller this distance, smaller the shadow gets. Currently,standard one-dimensional ASGs and FPDs are manufactured independently,and the 1D ASG is integrated on the protective cover of the FPD. Someoneexperienced in this Art is likely to follow the same approach tointegrate a 2D ASG with a FPD. As described in Example 4, such anapproach increases the distance between the top surface of the 2D ASGand the pixel array due to the air gap and protective cover of the FPD.As a result, 2D ASG's septa appear thicker in images than its physicalthickness. For example, 0.1 mm thick septa can appear as thick as 0.24mm in FPD images based on experiments.

If the 2D ASG is placed on the scintillator/pixel array, the distancebetween the top of the ASG and pixel array can be reduced by about 40%(the actual reduction depends on the properties of the FPD and 2d ASG).

One approach of the current invention also helps to reduce the change in2D ASG's shadow position due to gantry flex. Smaller the distancebetween top of the 2D ASG and pixel array, smaller the shift in shadowgets. As a result, grid shadow correction may be more robust.

B. The 2D ASG's septa not be aligned with detector pixel array and thegrid channel size is much larger than the detector pixel size. This way,2D ASG blocks less primary x-rays, and hence, improve image quality (SeeExample 5 and Example 6 for more details). One of the current inventionprototype 2D ASG demonstrates the benefits of this approach. In Example5 and Example 6, it is shown that the presently disclosed 2D ASGprovides 19% more primary x-ray intensity than a conventional 1D ASG.

This approach also brings a large flexibility to grid/detector design.Grid channel shapes do not need to be square, they can be hexagonal etc.(for example, hexagon shaped grid channels can produce smaller footprintthan square shaped grid channels). Also channel width (or grid pitch)can be varied across the detector (for example, grid height can be keptthe same channel width can be made smaller towards the central sectionof the detector where the scatter is highest. Channel widths can be madelarger at the periphery of the detector, where the scatter is lower.Variable grid channel width can reduce the cost of the 2D ASG).

C. To help to correct the 2D ASG's shadow, the current inventionutilizes pixelated scintillators rather than continuous scintillators inenergy integrating FPDs. Pixel walls are aligned with the septa of the2D ASG. As described in Example 5 and Example 6 and FIG. 22 , thisapproach reduces the long-range cross talk within the scintillator, andthus, reduces the intensity variations in 2D ASG's shadow.

Reduced cross talk also improves the spatial resolution to some extent.

A Method to Estimate and Correct Scatter Intensity Transmitted Through2D ASG.

Briefly, this method corrects the residual scatter intensity that wastransmitted through the 2D ASG. The method exploits the properties of 2DASG's shadow in FPD images, to estimate the residual scatter intensity.The details of the method are described in Example 7.

Besides correcting the residual scatter in images and improving imagequality, this method relaxes the technical requirements on 2D ASGfabrication and its potential implications on image quality.

For example, a 2D ASG with a given channel width, can be fabricated withsignificantly less height; aspect ratio of grid channels, known as gridratio, will be reduced which makes it technically less challenging andcheaper to fabricate. Moreover, lower grid ratio makes it easier tocorrect the grid shadows (i.e. the distance between the top of the ASGand pixel array reduced). While more residual scatter will reach the FPDdue to lower grid ratio, it can be “corrected” using the proposed methodin Example 7.

Example 2 Transmission Characteristics of a Two Dimensional AntiscatterGrid Prototype for CBCT [102]

Aim: High fraction of scattered radiation in CBCT imaging degrades CTnumber accuracy and visualization of low contrast objects. To suppressscatter in CBCT projections, a focused, two-dimensional antiscatter grid(2D ASG) prototype was developed. In this Example, the primary andscatter transmission characteristics of the 2D ASG prototype aimed forlinac mounted, offset detector geometry CBCT systems in radiationtherapy, are described and compared its performance to a conventionalone-dimensional ASG (1D ASG).

Methods: The 2D ASG is an array through-holes separated by 0.1 mm septathat was fabricated from tungsten using additive manufacturingtechniques. Through-holes' focusing geometry was designed for offsetdetector CBCT in Varian TrueBeam system. Two types of ASGs wereevaluated: a) a conventional 1D ASG with a grid ratio of 10, b) the 2DASG prototype with a grid ratio of 8.2. To assess the scattersuppression performance of both ASGs, Scatter-to-primary ratio (SPR) andscatter transmission fraction (T_(S)) were measured using the beam stopmethod. Scatter and primary intensities were modulated by varying thephantom thickness between 10 and 40 cm. Additionally, the effect of airgap and bow tie (BT) filter on SPR and T_(S) were evaluated. Averageprimary transmission fraction (T_(P)) and pixel specific primarytransmission were also measured for both ASGs. To assess the effect oftransmission characteristics on projection image signal-to-noise ratio(SNR), SNR improvement factor was calculated.

Results: In comparison to 1D ASG, 2D ASG reduced SPRs by a factor of 3to 6 across the range of phantom setups investigated. T_(S) values for1D and 2D ASGs were in the range of 21 to 29%, and 5 to 14%,respectively. 2D ASG continued to provide lower SPR and T_(S) atincreased air gap and with BT filter. T_(P) of 1D and 2D ASGs were 70.6%and 84.7%, respectively. Due to the septal shadow of the 2D ASG, itspixel specific primary transmission values varied between 32.5% and99.1%. With respect to 1D ASG, 2D ASG provided up to factor of 1.7 moreimprovement in SNR across the SPR range investigated.

Conclusions: When compared to a conventional 1D ASG, 2D ASG prototypeprovided noticeably lower SPR and T_(S) values, indicating its superiorscatter suppression performance. 2D ASG also provided 19% higher averageprimary transmission that was attributed to the absence of interseptalspacers and optimized grid geometry. The results indicate that thecombined effect of lower scatter and higher primary transmissionprovided by 2D ASG may potentially translate into more accurate CTnumbers and improved contrast resolution in CBCT images.

1. Introduction

High scattered radiation intensity is one of the major causes of imagequality degradation in FPD based CBCT, which leads to loss of contrastresolution, reduced CT number accuracy, and scatter induced imageartifacts [4]. Two major approaches, knows as scatter rejection andscatter correction methods, have been heavily investigated in the lastdecade to address this problem [5, 29, 36]. For scatter rejectionpurposes, ASGs developed for radiography and fluoroscopy have beenemployed in CBCT [27, 28, 39, 41, 42]. These ASGs consists of aone-dimensional array of radio-opaque, or at least radiation-absorbing,septa separated by aluminum or fiber spacers that support septa (suchASGs are referred as 1D ASG in the rest of the text).1D ASGs typicallyprovide a factor of 2 to 5 reduction in SPR values, and subsequentlyimprove CT number accuracy and reduce image artifacts. Moreover, bow tiefilters were also employed in CBCT, and they were shown to reducescatter fraction in sections of the object close to the CBCT isocenter[28, 72-75]. While such scatter suppression devices help to reducerelative scatter intensity, residual scatter reaching the FPD is stillhigh enough to deteriorate CBCT image quality. Scatter correctionmethods, which refers to correcting the effects scatter after itsdetection by the image receptor, are often employed to correct theresidual scatter [37, 38, 76-84]. Generally, both of these approacheshave been used together in CBCT systems for radiation therapy. However,the improvement in image quality is not at the desired level to achievehighly accurate CT numbers or improved visualization of low contrastobjects [21, 23, 65, 85].

To improve scatter suppression in CBCT, 2D ASGs may be viablealternative to 1D ASGs, since two-dimensional septa can potentiallyprovide better scatter rejection performance than one-dimensional septaemployed in 1D ASGs. With recent advances in advanced manufacturingmethods, 2D ASGs were introduced for mammography [86], breasttomosynthesis [87], and MDCT [88]. These 2D ASGs were fabricated usinglithographic techniques, and composed of copper (for mammography) ortungsten infused polymer (for tomosynthesis and MDCT) to achieve lowseptal thickness, high radio-opacity and geometric accuracy of the 2Dgrid. Moreover, due to the self-supporting structure of a 2D grid,interseptal spacers were eliminated in 2D ASGs, which may help toimprove 2D ASGs' primary transmission properties.

To assess the feasibility of 2D ASGs in the context of CBCT imaging, a2D ASG prototype was designed and fabricated to be employed in a linacmounted CBCT system. In contrast to 2D ASGs cited above, one embodimentof the current invention may be fabricated from pure tungsten usinglaser sintering based additive manufacturing methods. In this work, itsscatter and primary transmission characteristics under various imagingconditions were evaluated, and compared it with a standard 1D ASGemployed in the clinical CBCT system. Additionally, the impact oftransmission characteristics on the SNR improvement in projection imageswas assessed.

2. Materials and Methods 2.1. 2D ASG Prototype

The 2D ASG prototype was composed of a rectangular array of square,through-holes, separated by tungsten septa (FIG. 2 ). To minimize theshadow of septa in projection images, through-holes were aligned, orfocused, in two dimensions towards the x-ray source, and focusinggeometry was matched to TrueBeam's “half-fan” CBCT geometry (VarianMedical Systems, Palo Alto, Calif.). It was composed of two, 2 cm wideby 20 cm long modules, and they were glued together to achieve 2×40 cm²coverage on the FPD plane. 2D ASG has a grid pitch of 2.91 mm, gridheight of 23 mm, and a septal thickness of 0.1 mm, resulting in a gridratio of 8.2 (i.e. the ratio of grid height to through-hole width). 2DASG was fabricated using the Powder Bed Laser Melting (PBLM) process,and it was manufactured by Smit Röntgen (Best, Netherlands). PBLMprocess is similar to direct metal laser sintering, where a computerguided laser beam driven by the CAD model of the ASG traces and meltsthe tungsten powder that is placed on a built platform. The gridstructure is built on a layer-by-layer basis by lowering the platform,adding a new layer of tungsten powder, and repeating the laser meltingprocess.

2.2 Experimental Setup and Data Acquisition

Experiments were performed using the CBCT system in a Varian TrueBeamSTx linac. The CBCT system utilizes a PaxScan 4030CB FPD (Varian MedicalSystems, Palo Alto, Calif.) with an imaging area of 40×30 cm², and aGS-1542 x-ray tube (Varian Medical Systems, Palo Alto, Calif.). In allexperiments, half-fan CBCT geometry (also known as offset detectorgeometry) was utilized (FIG. 3 ), where the center of FPD was shifted by16 cm in the transverse direction with respect to imaging isocenter.Thus, the projected location of the beam's central axis (CAX) was 4 cmfrom the short edge of the FPD as indicated in FIG. 3 . Projectedlocation of CAX was simply referred as “CAX” in the rest of the text. Inall experiments, central 2×40 cm² section of the FPD was exposed tox-rays, and this section was referred as “measurement area” in the restof text. The remainder of the FPD was covered with 3.2 mm thick leadsheet that blocked 99.7% of the primary beam. TrueBeam CBCT system comeswith a focused, radiographic ASG (1D ASG), that was composed of a 1Darray of fiber interspaced lead septa. It has a grid ratio of 10, linerate of 60 l/cm, and septal thickness of 0.036 mm. In experiments withthe 2D ASG, 1D ASG was removed, and the 2D ASG was directly mounted onthe protective cover of the FPD. Three different ASG configurations wereevaluated: 1) Without ASG (NO ASG) 2) With 1D ASG, (i.e. the standardASG in TrueBeam CBCT) 3) With 2D ASG.

To modulate primary and scatter intensity, 30×30 cm² acrylic slabphantoms were employed, and they placed on the carbon-fiber treatmentcouch. The x-ray source was positioned at “0” degree gantry angle, suchthat the central axis of the x-ray beam was orthogonal to the couchsurface. As described further in Section 2.4, scatter intensity wasmeasured using lead beam stops, and they were placed between the phantomand the x-ray tube. To better visualize the spatial variations inscatter intensity, beam stop measurements were performed at 3 differentlocations (at CAX, 10 and 20 cm lateral to CAX) as indicated by redcircles in the detector plane.

All imaging experiments were performed at 125 kVp beam energy and with0.9 mm thick built-in titanium beam filter. The x-ray tube was operatedin pulsed mode, and tube current and pulse duration were adjusted toachieve sufficient signal intensity in images without saturating theFPD. The CBCT system comes with a built-in ion chamber placed on thex-ray tube exit window; the output of the ionization chamber was used tonormalize the projections due to changes in mAs settings and temporalvariations in tube output. The FPD was operated at 2×2 pixel binningmode (i.e. pixel size: 0.388 mm). For each measurement, 50 frames wereacquired, offset and flat field corrected, corrected for tube outputvariations, and averaged to reduce image noise. Corrected and averagedimages were used in extraction of primary and scatter intensities asdescribed in Sections 2.3 and 2.4. In FIG. 3, the top side view of theexperimental setup. Red circles correspond to measurement points, andthey are located at CAX, 10 cm, and 20 cm off-axis from CAX at detectorplane. In FIG. 3 , the bottom side beam eye view of the FPD. 2×40 cm²wide “measurement area” of the FPD is shown in gray. Rest of the FPD wascovered with Pb sheet.

2.3 Measurement of Primary Transmission

For primary transmission measurements, phantoms and the couch betweenthe x-ray source and FPD/ASG assembly were removed (Experiment setup 1in Table 1), and two image sets were acquired: one with and one withoutan ASG. The ratio of images acquired with and without ASG yielded theprimary transmission map. Average primary transmission fraction,(T_(P)), was obtained by averaging the values within a 1.6×1.6 cm²region of interest (ROI) in the primary transmission map,

Tp=(I(with ASG))/(I(without ASG))×100 Equation 1

where I is the average of pixel values within the predefined ROI. Thecenter of the ROI was centered across the short edge (2 cm) of themeasurement area, and shifted along the long edge (40 cm) to calculateT_(P) as a function of ROI location along the measurement area.

Since FPD pixels underneath the 2D ASG's septa receive lower intensityof primary x-rays, primary transmission varies spatially on apixel-by-pixel basis, which was not reflected in T_(P) values. Toevaluate this variation, primary transmission values extracted from theprimary transmission map were presented in cumulative histograms,referred as primary transmission histograms (PTH).

2.4 Measurement of Scatter Transmission Fraction and Scatter-to-PrimaryRatio (SPR)

For scatter intensity measurements, 3.2 mm thick, disc shaped Pb beamstops were mounted on a thin acrylic plate, between the phantom and thex-ray focal spot (FIGS. 4A&B). Beam stops attenuated 99.7% of theprimary beam, and the signal intensity in the beam stop shadow yieldedthe scatter intensity. In each beam stop shadow, a circular region ofinterest (ROI) was selected with a radius half of beam-stop shadow'sradius, and average scatter intensity, I_(S), was calculated byaveraging the pixel values within ROI. To keep the radii of beam stopshadows consistent across all scatter intensity measurements, the beamstop tray was placed at 85 cm from the x-ray focal spot. This way,magnification of beam stop shadows was kept constant at the detectorplane.

For Scatter-to-Primary ratio (SPR) measurements, two sets of images wereacquired at each experiment setup (Setups 2-4 in Table 1), one with andone without beam stops. SPR was calculated using following,

SPR=(I _(S))/(I _(P+S) −I _(S))  Equation 2

IS was obtained from images with beam stops as described above, whereasscatter plus primary intensity, IP+S, was obtained by averaging theimage signal at the same ROI location in images acquired without beamstops. To quantify the scatter suppression difference between the 1D and2D ASGs, SPR reduction factor was calculated, which was the ratio ofSPRs measured with 1D A and 2D ASGs.

Scatter Transmission fraction, T_(S), is the fraction of scatterintensity transmitted through an ASG, and it is a key metric inquantifying the scatter suppression performance of ASGs. T_(S) wascalculated using following,

T _(S)=(I _(S)(with ASG))/(I _(S)(without ASG))×100 Equation 3

Where IS was obtained from images acquired with and without ASG inplace.

As the beam stop size affects scatter intensity, SPR and T_(S)measurements were performed using 4 different diameter beam stops (theirdiameters were 3.5, 7.2, 10.5, and 13.6 mm at the detector plane), andthey were linearly extrapolated to “0” beam stop diameter by using leastsquares polynomial fitting [2, 89, 90]. In the Results section, SPR andT_(S) values at “0” beam stop diameter, and their standard errors werepresented. An example of measured SPR and T_(S) values versus beam stopdiameter was shown in FIGS. 4A&B. The data was measured using 20 and 40cm thick phantoms. Y-axis intercepts of linear fits yielded “0” beamstop diameter SPR and T_(S) values.

The effect of air gap on SPR and T_(S) was evaluated using 20 and 35 cmair gaps, and 20 cm thick phantom (Experiment setup 3 in Table 1). Airgap was the distance between the bottom surface of the treatment couchand the FPD plane. The effect of half-fan BT filter was evaluated using20 cm thick phantom and 20 cm air gap (Experiment setup 4 in Table 1).

TABLE 1 Experiment setups used during primary and scatter transmissionmeasurements Setups Phantom thickness (cm) Air gap (cm) BT filter Setup1 No object N/A No Setup 2 0-40 20 No Setup 3 20 20 and 35 No Setup 4 2020 Yes

In FIG. 4A SPR values for 20 and 40 cm thick phantoms were plotted as afunction of beam stop diameter. FIG. 4B T_(S) values for 20 and 40 cmthick phantoms were plotted as a function of beam stop diameter. Linearfits were used to extrapolate SPR and T_(S) values to “0” beam stopdiameter. Measurements were performed at 10 cm from CAX, and usingExperiment Setup 2 in Table 1.

2.5 Improvement in Signal to Noise Ratio

Improvement of low contrast object visualization in CBCT images is animportant subject as high scatter fraction degrades signal to noiseratio (SNR) in CBCT projections, and subsequently affects tissuevisualization in CBCT images [4]. The methodology to evaluate the effectof ASGs on SNR has been developed by several authors in the past [2, 44,91, 92], and it has been utilized in the context of CBCT [27, 42, 72].SNR for a contrast object is defined as

SNR=(cP)/(P+S)^(0.5)  Equation 4

where P is primary intensity, S is scatter intensity, and c is amultiplicative factor that defines the primary intensity difference, cP,between the contrast object and the uniform background [2]. Inevaluation of ASG's impact on SNR, SNR improvement rather than the SNRvalue by itself, has been assessed [44, 72]. Thus, SNR improvementfactor, K_(SNR), was employed in this evaluation, which is the ratio ofSNR with ASG to SNR without ASG. K_(SNR) was calculated as [2],

K _(SNR)=(T _(P)(1+SPR)^(0.5))/(T _(P) +T _(S)SPR)^(0.5)  Equation 5

where SPR was measured without ASG. K_(SNR) more than 1 indicates thatthe use of ASG increases SNR with respect to SNR without an ASG at agiven SPR value, whereas K_(SNR) less than 1 indicates that the ASGreduces the SNR. K_(SNR) for both ASGs were calculated from measuredSPR, T_(S), and T_(P) values described in Sections 2.3 and 2.4.

3. Results 3.1 Primary Transmission

FIG. 5A shows a section of the primary transmission map of the 2D ASG,where brighter and darker regions correspond to through-holes and septaof the 2D ASG, indicating higher and lower primary transmission values,respectively. The orange square shows an ROI that was used forcalculation of T_(P). T_(P) as a function of ROI location along themeasurement area is shown in FIG. 5B. The mean (and standard deviation)of T_(P) across all ROI locations were 70.6±0.2% and 84.7±0.4% for 1Dand 2D ASG, respectively. A slight reduction in 2D ASG's T_(P) isvisible at 20 cm. This location corresponds to the abutment surface ofthe two 2D ASG modules, where the septal thickness was doubled (i.e. 0.2mm). Thus, ROIs that included the location of the abutment surface hadrelatively lower T_(P) values.

To better quantify the variation in primary transmission, cumulativeprimary transmission histograms (PTH) were calculated FIG. 5C). For 1Dand 2D ASGs, minimum-maximum pixel specific primary transmission valueswere 68.3%-72% and 32.5%-99.1%, respectively. Although, FPD pixelsunderneath the 2D ASG's septa received low primary transmission, thepercentage of such pixels was relatively small; 97.2% of the pixelsreceived 50% or higher primary transmission (Arrow 1), and 75% of pixelsreceived 72% or higher primary transmission, which was maximum pixelspecific primary transmission value measured with 1D ASG. Finally, 57%of the pixels received 90% or more primary transmission with 2D ASG (redarrow 3). FIG. 5A shows a section of the primary transmission map of 2DASG, where bright and dark regions indicate higher and lower primarytransmission values, respectively. The orange colored square indicates a1.6×1.6 cm² ROI used for average primary transmission, T_(P),calculation. FIG. 5B shows T_(P) as a function of ROI location along thelong edge of the primary transmission map. FIG. 5C shows a cumulativeprimary transmission histograms of 1D and 2D ASGs.

3.2 Scatter to Primary Ratio (SPR)

SPR values as a function of phantom thickness and ASG configuration weremeasured using Experiment Setup 2 (Table 1), and results are shown inFIGS. 6A, 6C and 60C. As SPR values did not vary considerably acrossmeasurement points, only the results at 10 cm from CAX were summarizedbelow, and in Table 2. SPR without ASG increased from 1.11 to 8.44 as afunction of increasing phantom thickness, and 1D ASG reduced the SPRrange to 0.45-2.71. 2D ASG further reduced SPR range to 0.16-0.46. Whencompared to 1D ASG, 2D ASG provided a factor of 2.81 to 5.89 reductionin SPR, and SPR reduction by 2D ASG was more pronounced at largerphantom thicknesses.

The effect of air gap on SPR was evaluated using Experiment Setup 3(Table 1). SPR was reduced for all ASG configurations when the air gapwas increased from 20 cm to 35 cm (FIG. 7A). At 10 cm from CAX, SPRswith 1D and 2D ASG were reduced from 0.96 to 0.67, and from 0.27 to0.21, respectively (Table 3). At 35 cm air gap, 2D ASG provided a factorof 3.19 lower SPR than 1D ASG.

The effect of BT filter on SPR varied spatially in all ASGconfigurations (FIG. 7B): While BT filter reduced SPR within 10 cm ofCAX, SPR increased at 20 cm away from CAX for all ASG configurations. 2DASG continued to provide lower SPR values with respect to 1D ASG with BTfilter in place (Table 4). For example, at 10 cm from CAX, SPRs with 1Dand 2D ASGs were 0.68 and 0.18, respectively. SPR reduction factorsprovided by 2D ASG were 3.78 and 6.24 at 10 and 20 cm from CAX,respectively.

TABLE 2 SPR values as a function of phantom thickness at 10 cm from CAXpoint. Phantom SPR thickness SPR reduction (cm) NOASG 1D ASG 2D ASGfactor 10 1.11 ± 0.04 0.45 ± 0.01 0.16 ± 0.01 2.81 20 2.64 ± 0.09 0.96 ±0.02 0.27 ± 0.02 3.56 30 4.80 ± 0.20 1.54 ± 0.01 0.34 ± 0.03 4.53 408.44 ± 0.18 2.71 ± 0.03 0.46 ± 0.06 5.89

TABLE 3 SPR values as a function of air gap at 10 cm from CAX SPR SPRreduction Air gap (cm) NOASG 1D ASG 2D ASG factor 20 2.64 ± 0.09 0.96 ±0.02 0.27 ± 0.02 3.56 35 1.53 ± 0.02 0.67 ± 0.01 0.21 ± 0.01 3.19

TABLE 4 SPR values as a function of BT filter status at 10 and 20 cmfrom CAX. BT SPR Measurement filter SPR Reduction location status NOASG1D ASG 2D ASG factor 10 cm from No 2.64 ± 0.09 0.96 ± 0.02 0.27 ± 0.023.56 CAX 10 cm from Yes 1.82 ± 0.02 0.68 ± 0.02 0.18 ± 0.01 3.78 CAX 20cm from No 2.62 ± 0.07 1.04 ± 0.02 0.23 ± 0.01 4.52 CAX 20 cm from Yes 5.3 ± 0.12 1.81 ± 0.15 0.29 ± 0.01 6.24 CAX

3.3 Scatter Transmission Fraction (T_(S))

T_(S) as a function of phantom thickness was measured using ExperimentSetup 2, and results are shown in FIGS. 8A, 8B and 8C. At any givenphantom thickness and ASG configuration, T_(S) values varied less than2% across all measurement points. T_(S) for both ASGs were reduced asphantom thickness increased, that indicated better scatter suppressionperformance at large phantom thicknesses. At 10 cm from CAX, T_(S) of 1Dand 2D ASGs were in the range of 21.3-29.1% and 4.6-14%, respectively(Table 5). When compared to 1D ASG, reduction in T_(S) with 2D ASG wasmore pronounced at larger phantom thicknesses; when the phantomthickness increased from 10 to 40 cm.

As shown in FIG. 9A, T_(S) increased as air gap increased from 20 to 35cm for both ASGs, indicating that scatter suppression efficiency wasreduced at larger air gaps. At 10 cm from CAX, relative increase inT_(S) was 18% and 25% for 1D and 2D ASGs, respectively. 2D ASG stillprovided lower T_(S) with respect to 1D ASG (Table 6).

In contrast to SPR, T_(S) measured with BT filter did not exhibit largevariations across the three measurement points (FIG. 9B), and BT filterdid not make a large impact on T_(S); T_(S) increased slightly at CAX,and reduced monotonically further away from CAX with BT filter in place.T_(S) values measured with BT filter were within 1-3% of the valuesmeasured without BT filter. 2D ASG continued to provide lower T_(S)values with respect to 1D ASG (Table 7).

TABLE 5 T_(s) as a function of phantom thickness at 10 cm from CAX.T_(s) Phantom thickness (cm) 1D ASG 2D ASG 10 29.1 ± 0.1% 14.0 ± 0.7% 2025.6 ± 0.1%  8.8 ± 0.7% 30 22.2 ± 0.2%  5.7 ± 0.3% 40 21.3 ± 0.2%  4.6 ±0.3%

TABLE 6 T_(s) as a function of air gap at 10 cm from CAX. T_(s) Air gap(cm) 1D ASG 2D ASG 20 25.6 ± 0.1%  8.8 ± 0.7% 35 30.2 ± 0.3% 11.1 ± 0.2%

TABLE 7 T_(s) as a function of BT filter status at 10 and 20 cm fromCAX. Measurement T_(s) location BT filter status 1D ASG 2D ASG 10 cmfrom CAX No 25.6 ± 0.1% 8.8 ± 0.7% 10 cm from CAX Yes 26.2 ± 0.3% 8.9 ±0.6% 20 cm from CAX No 27.4 ± 0.1% 9.1 ± 0.1% 20 cm from CAX Yes 25.1 ±0.5% 5.9 ± 0.5%

3.4 SNR Improvement Factor

K_(SNR) was calculated for all measurement points, and all experimentsetups (See Table 1), and it was plotted as a function of SPR withoutASG (FIG. 10 ). FIG. 10 shows SNR improvement factors, K_(SNR), wereplotted as a function of SPR without ASG. K_(SNR) values were calculatedfor all measurement points and experiment setups (See Table 1). Cubicpolynomials (red and blue lines) were fitted to better visualize trendsin K_(SNR). Cubic polynomials were fitted to data points to visualizethe trends (blue and red lines). When compared to 1D ASG, 2D ASGprovided higher K_(SNR) values across the range of SPRs investigated inthis study, and SNR improvement provided by 2D ASG was more emphasizedparticularly at high SPR conditions. For example, at SPR of 8.7, K_(SNR)achieved by 1D ASG was 1.38, whereas 2D ASG provided K_(SNR) of 2.34,indicating a K_(SNR) increase by a factor of 1.7 with respect to 1D ASG.Change in air gap and presence of BT filter did not make any largedifference in K_(SNR) trends. When SPR was 1.1 or below, K_(SNR) of 1DASG was less than 1, indicating that SNR was degraded with respect toSNR without ASG. For 2D ASG, transition from SNR improvement to SNRdegradation occurred (i.e. K_(SNR) below 1) at SPR of 0.27. At SPR of 0,K_(SNR) of 1D and 2D ASGs were 0.84 and 0.92, respectively, whichimplies that degradation of SNR was less with 2D ASG in the absence ofscatter.

4. Summary and Discussions

A 2D ASG prototype aimed for CBCT systems in radiation therapy wasdeveloped and evaluated its x-ray transmission characteristics. As shownin Section 3.1, 2D ASG prototype provided about 20% higher primarytransmission on the average than the 1D ASG installed in the clinicalCBCT system. There are multiple factors that explain the higher averageprimary transmission provided by the 2D ASG. First, due to theself-supporting structure of the 2D grid, interseptal spacers are notneeded, whereas fiber spacers used between septa of 1D ASG thatattenuate the primary beam. The second factor is the effect of gridgeometry on primary transmission. The 2D ASG has a large grid pitch(2.91 mm) with respect to its septal thickness (0.1 mm) that leads to arelatively small footprint on the FPD38; the area covered by the 2DASG's tungsten septa constitutes less than 7% of the FPD's imaging area.On the other hand, 1D ASG has a relatively small pitch (0.167 mm) withrespect to its septal thickness (0.036 mm), and hence, its footprintcovers about 21% of the FPD's imaging area. Another factor that affectsprimary transmission is the suboptimal geometry and focusing of ASG'ssepta. Nonuniformities in septal thickness and deviations of septa fromideal focusing geometry would increase the shadow of septa inprojections, and subsequently, would reduce primary transmission.Geometric accuracy of 1D and 2D ASGs' septa cannot be measured directlyin projection images (due to the small septal thickness with respect toFPD pixel size), and therefore, its impact on primary transmission wasnot assessed in this study. However, spatial variations in ASG'sgeometric accuracy can be evaluated via observing spatial variations inT_(P). Standard deviation of T_(P) was 0.2% and 0.4% for 1D and 2D ASGs,respectively, while their mean T_(P) values were 70.6% (1D ASG) and84.7% (2D ASG). Such a small variation in T_(P) indicates that thegeometric accuracy of both ASGs was uniform across the measurement area.While 2D ASG provided higher “average” primary transmission, pixelspecific primary transmission varied between 32.5% and 99.1% due to the2D ASG's shadow in projections; FPD pixels that were underneath the 2DASG's septa received less primary beam with respect to pixels at thecenter of grid holes, as visualized in FIG. 5A. With 1D ASG, variationof pixel specific primary transmission was less than 4%, mainly due toits smaller septal pitch with respect to the FPD's pixel pitch.Although, 2D ASG provided lower primary transmission to a fraction ofpixels, the percentage of such pixels was relatively small, and 75% ofpixels still received more primary transmission than the maximum pixelspecific primary transmission provided by 1D ASG.

From scatter suppression point of view, 2D ASG provided significantlybetter performance in all imaging conditions that have been evaluated.When compared to the 1D ASG, SPRs measured were lower by a factor of 3-6with 2D ASG. Both ASGs provided better scatter suppression for largerobject thicknesses, as indicated by their lower T_(S) values at largerphantom thicknesses. When the air gap was increased from 20 to 35 cm,SPR was reduced in all ASG configurations. However, reduction in SPR wasdue to the relative reduction in scatter intensity at larger air gaps[27, 44, 93]. On the other hand, scatter rejection performance of bothASGs deteriorated at increased air gap, which was indicated by theincrease in T_(S) values at 35 cm air gap. This-observation was inagreement with other reports in the literature [41, 42]; as the air gapincreases, the angular distribution of scatter is less broad, andscattered radiation is less likely to be stopped by ASG's septa.

BT filter spatially modulates both primary and scatter beam intensity,and caused spatially nonuniform SPR distribution [28, 74, 94]. Whencompared to measurements without BT filter, SPR was lower in regionsclose to CAX where primary intensity was higher due to the thinner BTfilter section, and SPR increased further away from CAX, where primaryintensity was lower due to increased BT filter thickness. Theseobservations apply to both ASGs and NO ASG configuration, which arequalitatively in agreement with the literature [28, 75]. While BT filtercaused larger variations in SPR, its impact on the scatter suppressionperformance of 1D and 2D ASGs was less pronounced, as indicated by therelatively small difference (3.2% or less) in T_(S) values measured withand without BT filter.

Numerous studies have established that reduced scatter fraction in CBCTprojections translates into improved CT number accuracy and reducedimage artifacts [4, 27, 28, 41, 72, 95]. Thus, it is expected thatimproved scatter suppression performance of 2D ASG may be likely toincrease the CT number accuracy and reduce scatter-induced artifactswith respect to CBCT images acquired with 1D ASG. Besides improvement ofCT number accuracy, another area of interest is the improvement of lowcontrast resolution. While ASGs reduce scatter intensity, and have apositive effect on improving SNR, they also reduce primary intensity,that deteriorates SNR and contrast resolution. In addition to thetransmission characteristics of the ASG, relative intensity of scatter,or SPR, incident on the FPD determines the level of SNR improvement ordegradation with the use of ASG. Several studies have reported thatconventional 1D ASGs reduced SNR and contrast resolution in low tomedium scatter intensity environments, where SPR was below 1-2. In thisevaluation, SNR degradation with 1D ASG (i.e. K_(SNR)<1) occurred at SPRvalues below 1.1, and this result was in agreement with the literature[27, 28, 42]. Since SPR for various anatomical sites is generally lessthan two, the role of 1D ASGs in SNR improvement is typically limited tohigh SPR imaging conditions, such as CBCT imaging of Pelvis or abdomen[27, 28, 41, 42].

Across the SPR range investigated in this study, SNR improvement with 2DASG was up to a factor of 1.7 higher than 1D ASG. At higher SPR values,lower scatter transmission by 2D ASG plays an important role in SNRimprovement. At lower SPR values (e.g. SPR<1), higher primarytransmission by 2D ASG becomes a more dominant factor in SNRimprovement, as the scatter intensity constitutes a smaller faction ofthe total x-ray intensity incident on the FPD. The role of higherprimary transmission is particularly evident at “0” SPR condition, whereSNR improvement is solely determined by the primary transmissioncharacteristics of an ASG. At “0” SPR (FIG. 10 ), K_(SNR) of 1D and 2DASGs were 0.84 and 0.92, respectively. As a combined effect of bothlower scatter and higher primary transmission, 2D ASG provided SNRimprovement at SPR values down to 0.27. Therefore, the 2D ASG maypotentially improve contrast resolution in a larger range of SPRconditions with respect to conventional 1D ASGs.

One evaluation employed a 1D ASG with a grid ratio of 10, as it was theASG installed in the clinical TrueBeam CBCT system. While this is atypical grid ratio for commonly utilized 1D ASGs, 1D ASGs with gridratios above 15 have been developed in recent years. Such ASGs mayprovide improved scatter suppression and SNR performance in CBCTimaging. For example, Wiegert et al. [42] investigated a 1D ASG with agrid ratio of 27, and showed that T_(S) was a factor 2-3 lower withrespect to a 1D ASG with a grid ratio of 10. However, SNR improvementwith high grid ratio was worse, except in high scatter imagingconditions (e.g. SPR>5), which was attributed to poor primarytransmission characteristics of the ASG. Stankovic et al. [62] employeda 1D ASG with grid ratio of 21 in a linac mounted CBCT system. They haveshown that both CT number accuracy and contrast to noise ratio (CNR) wasimproved in a wide range of scatter conditions with respect to CBCTimages acquired without ASG. Fetterly and Schueler investigated asimilar 1D ASG for digital radiography [3]. They reported a factor 2-3lower T_(S) and comparable T_(P) with respect to 1D ASGs with moderategrid ratios, and better SNR improvement in a wide range of scatterconditions. Comparison of 2D ASGs with such high grid ratio 1D ASGs isan area remains to be investigated.

One of the challenges in implementation of 2D ASGs in CBCT is thecorrection of its septal shadow, or footprint, in projections. If notaddressed properly, spatial variations in image intensity may likely tolead to ring artifacts in reconstructed images. Moreover, variations inimage intensity may also lead to spatially nonuniform image noise inreconstructions. Similar challenges have been faced in utilization of 2DASGs in breast tomosynthesis [87], and nonlinear grid reciprocationschemes have been implemented to blur the ASG's footprint by moving theASG at high frequency during x-ray exposure [96]. A similar approach wasalso utilized for crosshatched ASGs in mammography [47]. Moreover, imagepost processing based correction algorithms were developed for 1D ASGsthat exploit the periodicity of septa footprint to suppress ASG'sfootprint in projections [97-99]. Such post processing approaches may aswell be implemented in the context of 2D ASGs.

1. Conclusions

A prototype 2D ASG was developed and its x-ray transmission propertieswere evaluated in half-fan geometry of a linac mounted CBCT system. Whencompared to a conventional 1D ASG, 2D ASG reduced SPRs by a factor of3-6, while providing 19% higher primary transmission on the average. Itwas observed that scatter suppression advantage of 2D ASG was maintainedat increased air gap and with bow tie filter in place. It is expectedthat lower SPR values achieved with 2D ASG may potentially translateinto reduced image artifacts and improved CT number accuracy in CBCT. Inaddition, when compared to 1D ASG, 2D ASG improved SNR in projections ina wide range of SPRs due to its higher primary transmission and scattersuppression capability. It is believed that improved SNR in projectionsmay lead to improved contrast resolution in CBCT images. While 2D ASGexhibited better scatter and primary transmission characteristics, itsseptal footprint leads to spatially varying primary transmission inprojections that may cause artifacts in reconstructed images. If thecorrection of its septal footprint is addressed, 2D ASG can be apromising scatter suppression device in improvement of CBCT imagequality in the future.

Example 3 Configuration of Current X-Ray Flat Panel Detectors and 1DASGS

Conventional, Potter-Bucky antiscatter grid: An array of 1D lead stripsand inter-spacers (so called Potter-Bucky grids). Such Potter-Buckygrids absorb scattered x-rays while transmitting primary x-rays. Suchgrids are an “add-on” component to the flat panel detector assembly.

Phosphor (scintillator layer): Phosphor layer absorbs incident x-raysand converts them to visible light photons. It is a continuous layerplaced (or directly deposited) on the detector's pixel array. Sincephosphor is a continuous layer, visible light photons from an x-rayinteraction, diffuse within the phosphor, and spread over to multiplepixels. Thus it reduces the spatial resolution of the detector.

Detector pixel array: It converts the visible light photons toelectrical charge, which is converted to digitized signal.

One of the disadvantages of this approach the increased distance betweenthe pixel array and the ASG, see FIG. 15 .

Example 4 Hybrid Flat Panel Detector of the Current Invention

The 1D Potter-Bucky grid described in Example 3 is replaced with oneembodiment of the current invention, a 2D antiscatter grid, and the 2Dgrid is directly placed on the phosphor layer such as in FIG. 16 . Acontinuous phosphor layer is replaced with a pixelated phosphor layer.Pixels in the phosphor are separated with reflective septa, preventingdiffusion of visible light photons neighboring detector pixels. The 2Dgrid's walls are aligned with the septa in the phosphor layer,minimizing the “inactive” area of the flat panel detector. This benefitcannot be achieved with the conventional approach previously described(i.e. the grid is mounted on the top of the protective detector cover;walls cannot be perfectly aligned with the septa of phosphor layer).

Some advantages of this hybrid design over existing flat panel detectorsinclude the following:

-   -   Air gap and detector cover between the 2D ASG and pixel array is        eliminated. As a result, the shadow of 2D ASG is minimized in        FPD images.    -   Pixelated phosphor structure reduces cross-talk between detector        pixels, improves spatial and contrast resolution (see FIG. 22 ).    -   2D grid provides better scatter absorption and improved primary        x-ray transmission with respect to conventional antiscatter        grids. In return, it reduces noise in CBCT images and improves        the accuracy of CT numbers (reduced noise leads to better soft        tissue visualization, improved CT number accuracy leads to        improved tissue density estimation in CBCT images).    -   Integration of pixelated phosphor with the 2D antiscatter grid        reduces the percentage of “inactive” detector area, thus more        primary x-rays will be detected by the detector. This approach        will reduce noise in CBCT images.

Although this approach may appear similar to the conventional CT scannerdetectors, it uses a flat detector and a different pixel arraytechnology. Flat detector employs either amorphous silicon or CMOS pixelarrays to reduce the pixel size and the footprint of the flat detector.

Example 5 Innovations in the Current Invention

Using a 2D grid and a pixelated scintillator in a flat panel detectorfor CBCT has not been previously described. The differences in grid andscintillator designs are important to achieve high image quality/costeffectiveness/fabrication feasibility. A viable design is both feasibleto fabricate and feasible to integrate with a flat detector. Moreimportantly, the design of the current invention should provide betterimage quality than competing grids and other technologies. As outlinedbelow, there are unique properties of the disclosed design that makesthe grid fabrication feasible and provide better image quality thanexisting technologies:

Adjustment of Focusing Geometry and Fabrication Method:

-   -   Focusing geometry should be adjusted such that grid's channels        are directed to the focal point of the x-ray source.    -   The method of grid fabrication can be critical to the        adjustability of the current invention to various system        configurations and geometries. In one embodiment, a laser        sintering process may be used for fabrication of the grids. In        one embodiment, the laser sintering process enables the        fabrication of the 2D grid with the desired physical        characteristics.

Using a Pixelated Scintillator in a Flat Panel Detector:

Pixelated scintillators are typically used to reduce the spread ofoptical photons within the scintillator, and to improve spatialresolution. Currently, pixelated scintillators are not used in flatdetectors, and there are good reasons for it: To improve the spatialresolution with a pixelated scintillator, the pixel size of thescintillator should match pixel size of the flat detector array. Flatdetector pixel size is about 200 microns (varies between 100 and 400microns, depending on the make and model). Thus, the pixel size of thescintillator should be in the order of 200 microns, which is challengingand expensive fabricate.

The pixelated scintillators of the current invention are primarily usedto improve the performance of 2D grids with a flat panel detector. Thepixel size of the scintillator is matched to the size of the 2D grid'schannels rather than the detector array pixels. As a result, the pixelsize of the scintillator will be 2-3 mm (feasible to fabricate). In thenext section, it is examined why this is a preferred approach. In a CTdetector, the detector pixel size is ˜1 mm (factor of 25 larger thanflat detector's pixel), and scintillators with 1 mm wide pixels can befabricated.

Selection of the Optimal Size of 2D Grid's Channels

In CT detectors, 2D grid's channels have the same size as the detectorarray pixels (channel size and the detector's pixel size are both ˜1mm). However, this approach is very challenging to implement in flatpanel detectors (˜200 micron detector pixel size): To match the flatdetector's pixel size, a 2D grid should have a 200 microns channelwidth. If the grid channel is 200 microns wide and channel walls are50-100 microns thick, a large percentage of the grid will be made ofchannel walls. Hence, most of the detector will be obstructed by thegrid's footprint, which would lead to very poor transparency for primaryx-rays. Since primary x-rays generate the useful signal in the detector,lower signal levels would lead to lower signal to noise ratio in images.

In the current invention, the approach is different than previousapproaches. The grid channel size is much larger than the detector pixelsize: the channel widths (or pitch) are in the order of 2-5 mms. Theselection of channel width depends on the channel wall thickness toachieve the desired primary x-ray transmission characteristics. The gridchannel width does not depend on the detector pixel size. For example,for a grid channel width of 3 mm, and a channel wall thickness of 100microns, the footprint of the grid shadow will be much less, and theprimary x-ray transparency will be higher. About 85% of the primaryx-rays will be transmitted through the grid, see FIG. 17 .

The 2D Grid's Channels does not Need to be Aligned with the DetectorPixel Array.

In a CT detector, 2D grid's channel walls are aligned with the deadspace between the CT detector's pixels. In fact, this proposed approachin flat detectors is described in US Patent application publication2004-0251420 A1 [100], herein incorporated by reference. However, atypical large area flat detector has up to 43 cm width. Alignment of 200micron wide grid channels with 200 micron wide detector pixels is quitechallenging over such a large length/area. This is a lesser issue in CTsystems as CT detectors are modular and smaller (each detector tile isless than 10×10 cm²).

In contrast to CT detectors and the patent application cited above,channel walls are not aligned with the detector pixels in the currentinvention. With this approach, the integration of the detector with the2D grid is more practical. Since the proposed 2D grid is not alignedwith detector pixels, any grid channel shape can be used in the 2D grid(rectangular, hexagonal etc.). As a consequence of my approach, pixelsin the shadow of the channel walls will have reduced signal. But, basedon my selection of the channel size (2-4 mm) and channel wall thickness(˜100 microns), such a signal reduction will not have a major impact inimage quality. (See FIG. 18 below). Only a small percentage of pixelswill be impacted by the shadow of the channel walls.

Selection of optimal wall thickness is important to minimize the effectof channel wall shadows on the image, and it is also crucial forfabrication feasibility.

To minimize the channel wall shadows in the image, the channel wallthickness should be as thin as possible. However, fabrication processbecomes more challenging for thinner wall thicknesses. One embodiment ofthe present invention, the laser sintering process is utilized, as wallthickness of 100 microns can be achieved and 2D grids up to 20×20 cm²can be fabricated. Thus, a 2D grid for a large area flat detector (e.g.40×40 cm²) can be built only from four 2D grid tiles. For wallthicknesses less than 100 microns, only two technologies exist(Vogtmeier et al. [88] and Tang et al. [101]. In one embodiment, thecurrent invention contemplates increasing the wall thickness to, say200-250 microns. Then other technologies can be used fabricate theproposed grid. However, if the wall thickness is increased to 200microns, the shadow of the channel walls may be too large, and detectorpixels in the wall shadows will receive little or no signal. Thesepixels will essentially be dead pixels, as shown in FIG. 18 . Insummary, in the current invention design (with 100 micron channel wallthickness and −3 mm wide channels), the signal reduction in the channelwall shadows will be at acceptable levels, and the fabricationfeasibility is preserved.

The 2D grid's shadow in the image must be corrected for high imagequality. This is a very important subject for the integration of 2Dgrids with flat detectors. If the shadow of the 2D grid is not fullycorrected, it will lead to image artifacts. If not addressed correctly,correction of grid shadows will be a major roadblock in implementationof 2D grids in flat panel detector systems.

Correction option 1: The magnitude of signal reduction due to grid'sshadow can be measured on a pixel by pixel basis under referenceconditions. This “calibration data” can be later used to process imagesand correct the reduction in signal due to grid. Such calibration databased correction methods are already known.

Correction option 2: The option 1 above may not be a robust correctionmethod due to long distance cross-talk in non-pixelated scintillators(see FIG. 21 ): Due to continuous (non-pixelated) scintillator layer inthe flat detectors, some of the light photons travel large distanceslaterally, and get detected by detector pixels far from the originallocation of x-ray absorption. This long distance cross-talk is alsoknown as glare. The magnitude of glare depends on the properties of theimaged object, therefore “the calibration data” based approach in Option1 cannot correctly handle the excess signal. The magnitude of veilingglare compromises the robustness of the correction option 1 above.

That's where the pixelated scintillators come into play (see thepixelated scintillator in FIG. 22 ). If the scintillator is pixelated,this long-distance cross talk will be prevented, and the reduced signalin the 2D grid's shadow can be better predicted. The prevention of longdistance cross-talk will also provide some improvement in spatialresolution.

To reduce long distance cross talk, scintillator's pixel size does notneed to be as small as detector's pixels (as in CT detectors). Thus, inthe disclosure, the pixel size of the scintillator will be matched tothe width of the grid's channels, and pixels of the scintillator will bealigned with channel walls of the 2D grid. With the alignment ofscintillator pixels and the grid's channels.

There are significant advantages to placing the 2D grid directly on thetop of the scintillator (rather than on the top of the protective coverof the flat detector). The first advantage is the ability to align 2Dgrid's channels with the pixelated scintillator, as described inpreviously. The other advantage is the reduced effective height of the2D grid (If the grid is installed on the protective cover, there will beabout a 1 cm gap between the grid and the scintillator layer). Reducedgrid height becomes important, when flat detector/grid assembly is usedin a CBCT gantry (Please see the “gantry flex and grid shadow”illustration in FIG. 20 ). A CBCT gantry typically “wobbles or flexes”slightly when it rotates around the patient due to flexing of the gantryarms under gravity. As a result, the position of the x-ray sourcechanges in relation to the grid/detector assembly, and the shadow of thegrid on the detector shifts. This “wobble or gantry flex” problem makesthe correction of the grid shadow challenging. Higher the grid the worsethe problem gets. Thus, by directly mounting the grid on thescintillator, the effective height of the grid can be reduced from −3 cmto −2 cm (physical height of the grid is assumed to be 2 cm).

Is the “gantry flex and grid height” an issue in CT systems? CT gantriesare built using a completely different architecture. They have minimalgantry flex/or wobble. Thus, the grid height is much less of an issue.

Example 6 Calculations of the 2D Grid's Physical Characteristics)

1) The slant or angle of the 2D grid's channels: Angle of the grid'schannels follow the divergence of the x-ray beam, thus the calculationof channel angles is not unique to the disclosed 2D grid design. Inother words, any grid design should incorporate similar channel anglessuch that channels are pointing towards the x-ray source.

2) Primary x-ray transmission characteristics: Primary transmission isprimarily a function of grid channel width and channel wall thickness.This is one of the fundamental properties of the 2D grid. Example 2describes the calculation of the primary transmission. FIG. 17 and FIG.18 are from Example 2. Essentially, the proposed grid providessignificantly better primary transmission that conventional antiscattergrids used with flat detectors (FIG. 17 ). Conventional grids transmitonly 60-70% of primary x-rays, whereas the 2D grid with 3 mm channelwidth and 100 micron wall thickness transmits 85% of primary x-rays(higher the primary transmission the better).

3) Scattered x-ray transmission characteristics: Grid channel's heightto width ratio (also known as grid ratio) is the primary factor thatdetermines the scatter transmission characteristics. Based onpreliminary experiments, a 2D grid with a grid ratio of 12 transmitsapproximately 6% of the scattered x-rays, whereas conventional grid witha similar grid ratio transmits 10-15% the scatter (lower the scattertransmission the better). Although theoretically 2D grid's grid ratio(and height) can be increased to provide even lower scattertransmission, it is not practically possible due multiple reasons(manufacturing challenges, thicker grid will be closer to the patient(patient hazard), makes it harder to correct grid's shadow due to gantryflex (as explained in item 5 above). In one embodiment, the grid ratioof a 2D grid will be around 6-15, and the height of the grid will be amaximum of 5 cm or so.). In one embodiment, a 2D grid with a grid ratioof 10, and a grid channel width of 3 mm, the grid height will be 3 cm.

4) Benefits of integrating pixelated scintillator and 2D grid with aflat panel detector array. Numerical calculations to predict itsbenefits are not straightforward, but can be done in the longer term(i.e. reduced long-range cross talk, improved corrections of 2D grid'sshadows). The integration of the 2D grid with the pixelated scintillatorprovides more benefits than the benefits of the individual components asit is a synergistic combination. A pixelated scintillator with 3 mmpixels can be incorporated into a flat detector without the 2D grid, andlong distance crosstalk can be reduced. However, the pixelatedscintillator and the reduced cross-talk also improves the correction of2D grid's shadow in the image.

Example 7 A Method to Correct Scatter Intensity Using a 2D AntiscatterGrid

Background: The purpose of the 2D antiscatter grid is to stop scatteredx-rays reaching the flat panel detector, and improve the image quality.However, a fraction of scattered x-rays can still pass through the 2Dgrid, and reach the detector. As a result, the image quality would bedeteriorated.

Although, 2D grid's height (or grid ratio) can be increased to reducethe transmission of scatter, such a grid design will lead to othertechnical and practical challenges. For example, increased grid heightis more difficult to fabricate, and it can also absorb more primaryx-rays (i.e. useful x-rays that form the image), which deterioratesimage quality.

Purpose of the innovation: The disclosed method utilizes the 2D griditself as a device to estimate the residual scatter transmitted throughthe 2D grid to the flat panel detector. Once the residual scatterintensity reaching the detector is estimated, the scatter intensity canbe “corrected” to improve image quality. This invention expands theutility of the 2D grid. In addition to being a scatter rejection device(as explained in my previous disclosure), the 2D grid is utilized as adevice to estimate and correct residual scatter reaching the detector.

How it works: The 2D grid's footprint introduces a unique pattern ofimage signal intensity variations as shown in FIG. 19 . In the absenceof scattered-rays (i.e. without any imaged object in the field), theratio of image intensity underneath the grid's footprint (red box 1) andat the center of a grid hole (red box 2) has a unique value. Whileimaging an object, majority of the scatter will be stopped by the 2Dgrid, and a small fraction of scatter will reach detector. As a result,this unique ratio of signal intensities in Boxes 1 and 2 will change.

If the ratio of signal intensities in Box 1 and 2 are measured withoutscattered x-rays, a calibration data can be generated. This calibrationdata can be used to estimate the scatter intensity when an object isimaged using the following equation.

S=I ₂((I ₁ /I ₂)−(P ₁ /P ₂))/(1−((P ₁ /P ₂)))  Equation 6

S=Scatter intensity in Box 2, when an object is imaged

I₁=Image intensity in Box 1 when an object is imaged

I₂=Image intensity in Box 2 when an object is imaged

P₁=Image intensity in Box 1 under calibration conditions (no scatter ispresent in the image)

P₂=Image intensity in Box 2 under calibration conditions (no scatter ispresent in the image)

The calculation described above can be repeated for any arbitrary “boxpairs” in the image shown in FIG. 19 . Thus, scatter intensity can becalculated for any point in the image to get a 2D scatter intensity map.Subsequently, scatter correction is achieved by subtracting this scatterintensity map from the image. The equation described above shows one wayof calculating the scatter intensity using the 2D grid's footprint inimages. Fourier domain methods may also be used to calculate the scatterintensity.

Thus, specific compositions and methods of a hybrid flat panel detectorfor cone beam CT systems have been disclosed. It should be apparent,however, to those skilled in the art that many more modificationsbesides those already described are possible without departing from theinventive concepts herein. Moreover, in interpreting the disclosure, allterms should be interpreted in the broadest possible manner consistentwith the context. In particular, the terms “comprise” and “comprising”should be interpreted as referring to elements, components, or steps ina non-exclusive manner, indicating that the referenced elements,components, or steps may be present, or utilized, or combined with otherelements, components, or steps that are not expressly referenced.

Although the invention has been described with reference to thesepreferred embodiments, other embodiments can achieve the same results.Variations and modifications of the present invention will be obvious tothose skilled in the art and it is intended to cover in the appendedclaims all such modifications and equivalents. The entire disclosures ofall applications, patents, and publications cited above, and of thecorresponding application are hereby incorporated by reference.

Example 8 A Computational Model of a 2D ASG Photon-Counting Detector

This example is directed to understand the physics of x-ray transmissionand detection in 2D ASG/photon counting detector assembly, andinvestigate the effect of various 2D ASG and detector configurations onCBCT image quality. The second goal of this example is to investigateaft fluence modulation approach, and identify optimal 2D ASG gridgeometries.

To study the x-ray transmission, we will employ an x-ray spectrum model[28], and GEANT4 for Tomographic Emission (GATE) Monte Carlo package tosimulate x-ray transmission through the phantom and 2D ASG [29]. Theimaging system geometry will mimic the imaging conditions of a linacmounted CBCT system, and we will employ various phantom setups depictingvarious anatomical sites. We will also simulate the CdTe detector tostudy the effects detector backscatter. GATE simulations will providethe incident x-ray spectrum on the detector, and we will investigate how2D ASG geometry (such as grid height, pitch, septal thickness, and septafooting) affects the transmitted spectrum of primary and scatteredx-rays. We will also study the characteristic x-rays emitted fromtungsten septa. While the overall contribution of tungstencharacteristic x-rays to spectrum contamination is expected to be small,it may have a larger contribution in pixels located in septal shadows.As, our proposed Aft Fluence Method employs pixel counts in septalshadows, we will evaluate whether spectral contamination in septalshadows affect the efficacy of our fluence modulation approach.

To investigate the detected energy spectrum in the CdTe detector, wewill model the detector response, or signal formation, by adopting themethodology used by Schmidt [30]. Modeling the detector's energyresponse with GATE simulated incident x-ray spectrum will enable us tostudy the effects of charge sharing and the “detected” energy spectrumfor a variety 2D ASG geometries and pixel sizes.

Example 9 Experimental Evaluation of 2D ASG Scattered RadiationReduction

We will measure scatter and primary transmission characteristics of the2D ASG, using well-established approaches [17], and determine whethersufficient spectral decontamination is achieved by performing dualenergy CBCT imaging experiments; we will assess the noisecharacteristics of CBCT images, contrast-to-noise ratio (CNR)improvement, and accuracy of iodinated contrast quantification in avariety of phantoms mimicking clinically relevant anatomical shapes. Wewill also be able to quantify the spatial variations in spectralcontamination (i.e. spectral contamination in septal shadows versuscenter of grid holes). These studies will be performed at low x-rayfluences to minimize the effect of pulse-pile up on our evaluations.

We will setup a benchtop CBCT system using a fixed anode x-ray tube, alinear photon counting detector, and a rotation stage for the phantom.As such linear detectors are about 40-50 mms in length, we will performmultiple acquisitions for each projection by translating the detector inthe transverse plane.

While a linear detector will allow us to reconstruct only the centralCBCT slice, we will emulate the scatter conditions of full CBCT exposuregeometry (the size of the exposed area will be about 30×40 cm2 atdetector plane). To suppress potential grid artifacts caused by septalshadows, we will explore various detector calibration approaches, andimplement a total variation minimization (TVM) based grid artifactsuppression algorithm.

Example 10 AFT Fluence Modulation Validation

This example characterizes improvement in count rate and energyresolution performance under clinically relevant fluence environments.

We will test two 2D ASG septal configurations one with standard(constant thickness) septa and the other one with septal footing. First,we will characterize the pixel response as a function of incidentfluence, and assess the effects of pulse pile-up both in septal shadowsand in the center of the through-holes; we will observe the differentialchanges in count rate changes in these two regions and determine theconditions that will employ the Aft Fluence Method. As the detector canbe translated with respect to a fixed 2D ASG in our benchtop system, wewill be able to change the grid position with respect to detector pixelsin a precise manner, and investigate the effects of grid/pixel alignmenton Fluence modulation patterns and pixel-specific energy resolution.

Specifically, we will evaluate the energy response of pixels neighboringthe septal shadows; as we hypothesize that energy resolution in suchpixels can be improved due to reduced charge sharing (described inSection 2.2). We will assess the magnitude of improvement in energyresolution in such pixels. If we can successfully demonstrate thatenergy resolution is improved in such pixels, novel photon countingdetectors can be developed in the feature, where grid septal pitch ismatched to detector pixel pitch, reducing charge sharing among allneighboring pixels, and hence, improving energy resolution of photoncounting detectors.

We will also perform an image quality comparison of photon-counting CBCTto single energy FPD based CBCT. This step is aimed to demonstrate theutility of proposed solutions as well as the benefits of photon countingapproach in CBCT under realistic imaging conditions.

Example 11

The present technology is pertinent to x-ray imaging systems that employantiscatter grids (2D or 1D). As described in this Example 11, a methodmay be provided to realign a 2D or 1D grid when the detector is offsetwith respect to the x-ray source, or vice versa. Detector (or x-raysource) offsetting may be particularly important in the context of CBCTsystems that employ flat panel detectors. With this method, the sameantiscatter grid can be used in both centered and offset detectorgeometry.

FIG. 32 presents a diagram of a centered detector CBCT geometry for aCBCT system 3200 according to a known embodiment. In a typical CBCTsystem 3200, an x-ray source 3210 and an x-ray detector 3220 rotatesaround the center 3230 of rotation 3240 to generate a 3D image. A fieldof view (FOV) 3250 is indicated by the circular shaded region in FIG. 32. In the known system 3200, to make the FOV 3250 larger, a largerdetector 3220 is needed. However, this may not be feasible in mostinstances, as such large detectors may not be available, or they can beexpensive. In some cases, as where the x-ray source 3210 and detector3220 are situated in an O-shaped gantry 3260, sizing considerations maycome into play such that it is not possible to use a larger detector3220 given limitations on dimensions of the O-shaped gantry 3260.

FIG. 33 presents a diagram of an offset detector CBCT geometry for theknown CBCT system 3200 shown in FIG. 32 . With further reference to FIG.32 , the known technique shown in FIG. 33 may make the FOV 3250 largerusing the illustrated offset detector geometry. In this known technique,the detector 3220 is shifted laterally, and only half of the FOV 3250 isimaged during each detector 3220 position. Information about the fullFOV 3250 is collected when the detector 3220-source 3210 system rotates360 degrees. Although this technique using offset detector geometry iswell known for a CBCT system such as system 3200, the ability to useboth centered and offset detector geometry in the same CBCT system onlyworks when there is no 2D antiscatter grid present on the detector 3220side.

FIG. 34A again presents the 2D ASG of FIG. 1 . In the illustratedembodiment, the 2D antiscatter grid (ASG) 3410 according to the presenttechnology is on a side of the x-ray detector 3220 that faces the x-raysource 3210 during use. In some embodiments, the 2D ASG 3410 accordingto the present technology may be a 2D array of through-holes separatedby metal, preferably tungsten, septa. The 2D ASG 3410 may be placed onthe detector 3210 to reduce scattered x-ray radiation and therebyimprove image quality.

FIG. 34B presents a schematic diagram of an x-ray imaging system 3400according to some embodiments of the present technology. As used in CBCTfor example, each through-hole of the 2D ASG 3410 may need to be alignedto angularly face the x-ray source 3210 so that x-rays emitted by thesource 3210 can go through the holes of the 2D ASG 3410. Thus, inpractice, particular manufactures of the 2D ASG 3410 according to thepresent technology may need to be customized based on a particulardetector 3220-x-ray source 3210 geometry. In the example shown in FIG.34B, the 2D ASG 3410 is tailored for the above described centereddetector CBCT geometry. Accordingly, the through-holes of the 2D ASG3410 are designed to point towards the x-ray source 3210 to match thedivergence of the x-rays emitted from the source 4320 during use.However, if the detector 3220 is shifted laterally to increase the FOV3250, as in the offset detector CBCT geometry and associated techniquesdescribed above with reference to FIG. 33 , the 2D ASG 3410 will nolonger be aligned towards the x-ray source 3210.

FIG. 35 presents the system 3400 shown in FIG. 34B where the 2D ASG 3410according to the present technology is shifted laterally. When thedetector is shifted laterally, the 2D ASG 3410 would be focused towardsa point 3510 in space where the x-ray source 3210 is not present. Giventhe lack of full alignment, x-rays emitted by the source 3210 will notgo through the holes of the 2D ASG 3410 as desired and since not allx-ray radiation will reach the detector 3210 the image quality willsuffer. Thus, if a CBCT system like system 3400 uses both centered andoffset detector geometries, it will need two different 2D grids—onedesigned for centered detector, and another one for offset detectorgeometry. Having separate 2D grids for each geometry can presentchallenges in practice. The method described in this Example 11 providesa technique to address this challenge.

FIG. 36 presents an x-ray imaging system 3600 according to someembodiments of the present technology. In some embodiments, system 3600may be a CBCT imaging system. The technique depicted in FIG. 36 startswith the 2D ASG 3410 designed for centered detector geometry, as shownand described above. Therefore, centered detector CBCT imaging can beperformed without encountering the aforementioned technical challenges.When the same single 2D ASG 3410 is to be used with offset detectorCBCT, the method according to the present technology may utilize thelateral shifting along with a tilting of the detector 3220. This tiltingmoves the detector 3220 closer to the x-ray source 3220, such that thesepta (and/or through-holes thereof) of the 2D ASG 3410 are all stillfocused towards the x-ray source 3210.

If the detector 3220 is shifted laterally by a first distance (denotedas “x” in FIG. 36 ) for offset geometry, the detector 3220 needs to betilted by an angle (denoted as theta “θ” in FIG. 36 ) so that a shiftingby a second distance (denoted as “y” in FIG. 36 ) provides the alignmentof the through-holes of the 2D ASG 3410 toward the x-ray source 3210. Inthe embodiment of system 3600 shown in FIG. 36 , one end of the detector3220 defines the tilting origin 3610 for angle θ, the first distance (x)is defined between a center 3620 of the detector 3220 and a normal line3630 from the center 3620 to the x-ray source of the initial centereddetector CBCT configuration, and the second distance (y) is definedbetween a reference plane 3640 and the center 3620, with plane 3640being orthogonal to line 3630.

FIG. 37A presents a schematic diagram providing additional details forthe shifting and tilting technique described above for the system 3600of FIG. 36 . With the system 3600 in the initial centered detector CBCTconfiguration (e.g., 3600-1), a length of the normal line 3630 betweenthe x-ray source 3210 and the center 3620 of the detector 3220 isdenoted as “SDD” (source to detector distance) in FIG. 37A. SDD is theshortest distance on the virtual line 3620 that is connecting the x-raysource 3210 to the detector 3220. This normal line 3630 is alsoorthogonal to the detector entrance surface. From the initial state3600-1, the detector 3220 may be laterally shifted by the amount x totake on the state labeled 3600-2 in FIG. 37A. As discussed above withreference to FIG. 36 , the septa (and/or through-holes thereof) of the2D ASG 3410 are not fully aligned with the x-rays emitted by the source3210 in the state 3600-2. In the technique according to the presenttechnology, a third state 3600-3 of the system 3600 may include thedetector 3220 being tilted by the angle θ, which results in the center3620 of the detector 3220 being shifted upward by the distance y.

The values of angle θ and the distance y are functions of the SDD andthe value of x, as shown in equations (1) and (2), below:

$\begin{matrix}{{\theta\left( {{tilt}{angle}{of}{detector}3220} \right)} = {\arcsin\frac{x}{SDD}}} & (1)\end{matrix}$y(upward movement of the center of the detector)=SDD−√{square root over(SDD²−x²)}  (2)

As a result of the tiling by angle θ after the lateral shifting, thesystem in its state 3600-3 has the septa (and/or through-holes thereof)of the 2ASG 3410 of the detector 3220 fully aligned with the x-raysemitted by the source 3210 during operation of system 3600.

FIG. 37B presents a flowchart of a method 3700 to improve (or assure)alignment of an antiscatter grid when the detector is shifted withrespect to the x-ray source, according to some embodiments of thepresent technology. With further reference being made to FIGS. 35, 36and 37A, method 3700 may include the step of arranging 3710 x-ray source3210 and detector 3220 having an ASG in the above described centeredgeometry. In some embodiments, detector 3220 as used in method 3700 maybe a flat panel detector. In some embodiments, the ASG of the detector3220 may be a 2D ASG (e.g., 3410) as described herein according to thepresent technology. In other embodiments, the ASG of detector 3220 maybe a 1D ASG. Notably, method 3700 may be performed for the advantageoustechnical benefits as described herein in either the case of a 1D ASG ora 2D ASG.

In the centered geometry, through-holes of the ASG may be aligned withx-ray emission paths of the x-ray source 3210. Method 3700 may includethe step of positioning 3720 an object between detector 3220 and x-raysource 3210. In one embodiment, the positioning 3720 step of method 3700may be performed after the arranging 3710 step. In other embodiments,the positioning 3720 step of method 3700 may be performed before thearranging 3710 step.

Method 3700 may include the step of imaging 3730 the object in a firstFOV with the detector 3220 and x-ray source 3210 arranged 3710 in thecentered geometry. With the object imaged 3730 in the first FOV, method3700 may include the step of arranging 3740 the x-ray source 3210 andthe detector 3220 in the above described offset geometry. In the offsetgeometry, through-holes of the ASG may be at least partially unalignedwith the x-ray emission paths of the x-ray source 3210. With the source3210 and detector 3220 arranged 3740 in the offset geometry, method 3700may include the step of moving 3750 the detector 3220 to realign thethrough-holes of the ASG with the x-ray emission paths of the x-raysource 3210. With the detector 3220 so moved 3750, method 3700 mayinclude the step of imaging 3760 the object in at least a second FOV.

In some embodiments, the arranging 3740 step of method 3700 may includeshifting 3765 the detector 3220 laterally by the first distance x, asdescribed above with reference to FIGS. 35, 36 and 37A. The laterallyshifting 3765 step of method 3700 may include positioning 3770 one end(e.g., at the point labeled 3610) of the detector 3220 proximal to anintersection of the normal line 3630 defined from the x-ray source 3210and the reference plane 3640 of the detector 3220 and having a lengthSSD, as shown in FIGS. 36 and 37A.

In one embodiment of method 3700, the moving 3750 of method 3700includes tilting 3775 the detector 3220 upwardly toward the x-ray source3210 by the angle θ, where θ has its origin 3610 at the aforementionedone end of the detector 3220. In the embodiment, the moving 3750 ofmethod 3700 also includes upwardly shifting 3780, e.g., as a result ofthe tilting 3775, the center 3620 of the detector 3220 by the seconddistance y, as described above with reference to FIGS. 36 and 37A.

Example 12

FIGS. 38A-38D present an x-ray imaging system 3800 utilizing a flatpanel detector (FPD) 3810 with a 2D ASG (e.g., 3410 as shown in FIG.34A), according to some embodiments of the present technology. In theillustrated embodiments of FIG. 38A-38D, the FPD 3810 may be embodiedin, for example and without limitation, a 40×30 cm FPD. In someembodiments, system 3800 may be a CBCT imaging system. With furtherreference being made to FIGS. 34A-37B, system 3800 may include FPD 3810and an x-ray source 3820. X-ray source 3820 may include one or more bowtie filters 3860 (e.g., BT1, BT2 and/or BT3). When energized, x-raysource 3820 may emit x-ray radiation according to an x-ray emission path3850. System 3800 may thus be employed for generating x-ray images of anobject 3840 positioned in the x-ray emission path 3850.

System 3800 may include a rotation and translation stage 3830. In someembodiments, the rotation and translation stage may include the FPD 3810with its associated 2D ASG 3410. The FPD 3810 may define theabove-described reference plane 3640. FIG. 38B depicts system 3800 withthe FPD 3810 and x-ray source 3820 positioned with respect to oneanother in the above-described centered geometry. FIG. 38C depictssystem 3800 with the FPD 3810 and x-ray source 3820 positioned withrespect to one another in the above-described offset geometry.

Referring now to FIGS. 38B-38D, the rotation and translation stage 3830may include rotation means 3870 operably coupled to the FPD 3810. Forexample and without limitation, the rotation means 3870 may include, orbe embodied in, a motor and associated mechanism (e.g., gears) operablycoupled to a portion of FPD 3810 (e.g., proximal to the center 3620 ofthe FPD 3810), and configured to, or otherwise capable of, rotating theFPD 3810 in a clockwise 3875 and/or counter-clockwise 3880 directionrelative to an axis z, as shown in FIG. 38B. In examples of the presenttechnology where rotation means 3870 includes, or is embodied in, amotor, means 3870 may be further operably coupled (e.g., by a wiredelectrical connection) to a motor controller and/or driver 3885 that canregulate a magnitude and/or polarity of an electric current to means3870 to thereby control a speed and/or direction of rotation of a motorshaft thereof.

In some embodiments, the rotation and translation stage 3830 may includelateral shifting means 3890 operably coupled to the FPD 3810. Forexample and without limitation, the lateral shifting means 3890 mayinclude, or be embodied in, a motor and associated mechanism (e.g.,gears) operably coupled to a portion of FPD 3810, and configured to, orotherwise capable of, laterally shifting the FPD 3810 in alternatedirections along an axis x, as shown in FIG. 38C. The embodiment shownin FIG. 38C illustrates a resulting lateral shifting using means 3890 bythe first distance x as between the FPD 3810 center 3620 and the normalline 3630. In examples of the present technology where lateral shiftingmeans 3890 includes, or is embodied in, a motor, means 3890 may befurther operably coupled (e.g., by a wired electrical connection) to themotor controller and/or driver 3885 that can regulate a magnitude and/orpolarity of an electric current to means 3890 to thereby control a speedand/or direction of rotation of a motor shaft thereof.

In some embodiments, the rotation and translation stage 3830 may includeupwardly shifting means 3895 operably coupled to the FPD 3810. Forexample and without limitation, the upwardly shifting means 3895 mayinclude, or be embodied in, a motor and associated mechanism (e.g.,linear actuator 3897) operably coupled to a portion of FPD 3810 (e.g.,proximal to the center 3620 of the FPD 3810), and configured to, orotherwise capable of, upwardly shifting the center 3620 of FPD 3810 inalternate directions (e.g., upward and downward) along an axis y, asshown in FIG. 38D. In examples of the present technology where upwardlyshifting means 3895 includes, or is embodied in, a motor, means 3895 maybe further operably coupled (e.g., by a wired electrical connection) tothe motor controller and/or driver 3885 that can regulate a magnitudeand/or polarity of an electric current to means 3895 to thereby controla speed and/or direction of rotation of a motor shaft thereof. In someembodiments, upward shifting of FPD 3810 by the second distance y usingmeans 3895 results in tilting FPD 3810 upwardly by the aforementionedangle θ with respect to reference plane 3640.

In some embodiments of the present technology, the system 3800 asdiscussed above in this Example 12 may be utilized to, at least in part,implement or otherwise perform one or steps of the method 3700 asdescribed above in Example 11.

Example 13

The present technology is also pertinent to portable or compact x-rayimaging systems that employ the disclosed 2D ASG (e.g., 3410). Asdescribed in this Example 13, a portable or compact CT scanner andassociated methods may be provided that provides superior imagingquality as compared to known portable or compact CT scanners.

FIGS. 39A-39F present a compact or portable CT scanner system 3900presents utilizing an FPD 3920 with a 2D ASG as shown in FIG. 34A (e.g.,2D ASG 3410), according to some embodiments of the present technology.CT scanner system 3900 includes an x-ray source 3910 positioned on andoperably coupled to an O-shaped gantry 3930. System 3900 also includesthe FPD 3920 positioned on and operably coupled to the O-shaped gantry3930. The O-shaped gantry 3930 defines an opening 3945 having a centeraxis 3980 into which an object 3940 to be imaged may be positioned. FPD3920 includes the 2D ASG 3410 which faces the opening 3945. In someembodiments, a diameter of the opening 3945 may be 35-45 cm, suitablefor imaging of head and neck region, and extremities, or human patients.To scan the human torso, bore diameter can be increased.

In some embodiments, the x-ray source 3910 is positioned radiallyopposite the FPD 3920 on the O-shaped gantry 3930. System 3900 mayinclude a controller or computer 3990 operably coupled to x-ray source3910 and/or FPD 3920. System 3900 may also include means (not shown inFIGS. 39A-39F) for rotating the O-shaped gantry 3930 in a clockwise 3960and/or counterclockwise 3970, annular (e.g., circumferential) direction.For example and without limitation, the means for rotating the O-shapedgantry 3930 may include, or be embodied in, a motor and associatedmechanism (e.g., gear(s) or belt(s)) operably coupled to a portion ofO-shaped gantry 3930 (e.g., proximal to the outer circumference ofO-shaped gantry 3930), and configured to, or otherwise capable of,causing the entire O-shaped gantry to rotate about the center axis 3980in alternate directions (e.g., clockwise and counterclockwise). Inexamples of the present technology where the means for rotating theO-shaped gantry 3930 includes, or is embodied in, a motor, such meansmay be further operably coupled (e.g., by a wired electrical connection)to the controller or computer 3990 that can regulate a magnitude and/orpolarity of an electric current to the means to thereby control a speedand/or direction of rotation of a motor shaft thereof. In someembodiments, rotation of the O-shaped gantry 3930 causes the radialalignment of the x-ray source 3910 and the FPD 3920 to be maintained atall times during the rotation such that an x-ray emission path 3950 fromsource 3910 to FPD 3920 beneficially passes through all the septa(and/or through holes thereof) of 2D ASG 3410, as described above inExample 12, for instance.

The O-shaped gantry 3930 may be positioned in a correspondingly shapedhousing 3955. The O-shaped gantry 3930 may include a pair of spindles3975 positioned radially opposite one another. In some embodiments,system 3900 may be an assembly of a base 3925, two stanchions 3915 and aset of wheels 3935 coupled to the base 3925. In one example, spindles3975 may extend through the housing 3955 and be rotatably coupled to thestanchions 3915. In another example, the housing 3955, but not also theO-shaped gantry 3930, includes the spindles 3975.

System 3900 may also include means (not shown in FIGS. 39A-39F) forrotating the O-shaped gantry 3930 and/or housing 3955 in a clockwise3965 and/or counterclockwise 3985, radial direction. For example andwithout limitation, the means for rotating the O-shaped gantry 3930and/or housing 3955 may include, or be embodied in, a motor andassociated mechanism (e.g., gear(s) or belt(s)) operably coupled to aportion of O-shaped gantry 3930 and/or housing 3955 (e.g., proximal tothe outer circumference of O-shaped gantry 3930 and/or housing 3955),and configured to, or otherwise capable of, causing the entire O-shapedgantry and/or housing 3955 to rotate about a radial axis 3905 inalternate directions (e.g., clockwise and counterclockwise).

In examples of the present technology where the means for rotating theO-shaped gantry 3930 and/or housing 3955 includes, or is embodied in, amotor, such means may be further operably coupled (e.g., by a wiredelectrical connection) to the controller or computer 3990 that canregulate a magnitude and/or polarity of an electric current to the meansto thereby control a speed and/or direction of rotation of a motor shaftthereof. In some embodiments, rotation of the O-shaped gantry 3930causes the radial alignment of the x-ray source 3910 and the FPD 3920 tobe maintained at all times during the rotation such that an x-rayemission path 3950 from source 3910 to FPD 3920 beneficially passesthrough all the septa (and/or through-holes thereof) of 2D ASG 3410, asdescribed above in Example 12, for instance.

The CT scanner system 3900 according to the present technology isportable and may also be fabricated in an ultracompact design given thatFPD 3920 with the 2D ASG 3410 is included. Embodiments of the presenttechnology that provide for rotation of the O-shaped gantry and thusalso the x-ray source 3910 and the FPD 3920 about radial axis 3905represents a substantial advance in the state of the art as knownportable CT scanners do not have this gantry tilting capability.

Scattered radiation is a major cause of image quality degradation inknown flat panel detector-based CT scanner. The use of the 2D ASG 3410according to the present technology addresses this problem, and allowsthe use of FPD 3920 in CT for soft tissue imaging. The wheels of system3900 enable it to be transported from one room to another. Suitablepower supply electronics may be positioned in the base 3925, forinstance. In some embodiments, system 3900 according to the presenttechnology may be fitted into land, sea, or air vehicles for use byfirst responders. The controller 3990 may further include, or beoperably coupled to, communications and networking components 3992, suchthat imaging data may be transmitted to clinician workstations orcomputing devices 3996 positioned remote from the system 3900 for suchadvantageous technical effects as enabling telemedicine support via acommunications network 3998 (e.g., Internet, satellite, cellular, etc.).

The portable and, in some embodiment, ultracompact CT scanner system3900 according to the present technology may be employed for x-rayimaging of extremities, the head, and/or back of patients. Thedimensions of system 3900 and its various component parts may be scaledup to provide CT scanning capability for the human torso. The FPD 3920of system 3900 may include, or be embodied in, an energy integrating orphoton counting detector. In some embodiments, and area of the FPD 3920may be 20 cm×30 cm. In other embodiments, the area of FPD 3920 may befrom 20 cm×20 cm to 43 cm×43 cm.

In some embodiments, dimensions of FPD 3920 may be beneficiallyminimized to, for example and without limitation, thereby enable a morecompact CT scanner system 3900 by utilizing the offset detectorgeometry, such as is described above in Example 12. The offset detectorgeometry may allow for coverage of a larger FOV with a smaller FPD 3920.

In some embodiments, the CT scanner system 3900 according to the presenttechnology may be used to implement, at least in part, any or all of thesteps of the method 3700 as described above with reference to FIG. 37Bin Example 12, for instance. In such example implementations, apositioning technique such as the method 3700 may enable switching theFPD 3920 and thus also the 2D ASG 3410 of system 3900 from the centereddetector geometry to the offset detector geometry to thereby enable morecompact dimensions and variable imaging applications and FOVs dependingupon a specific clinical need at hand.

In some embodiments of system 3900, a separation distance (e.g., pitch)between two adjacent vertical walls of the 2D ASG 3410 may be from 0.8mm to 3 mm. A grid ratio of the 2D ASG 3920 in system 3900 may be from 4to 16. The pitch of the 2D ASG 3410 can be different from a detectorpixel pitch of the FPD 3920. For example, and without limitation, thepitch of the 2D ASG 3410 may be 2 mm and the detector pixel pitch of FPD3920 may be 0.35 mm. In another embodiment, the pitch of the 2D ASG 3410may be 2 mm and the detector pixel pitch of FPD 3920 may be 0.2 mm.

In some embodiments, the grid pitch of 2D ASG 3410 may be larger thanthe detector pixel pitch of FPD 3920. In system 3900, the vertical wallsof 2D ASG 3410 may be formed at least in part of a radiation-opaque, orat least radiation-absorbing, material such as a metal. As installed inembodiments of the CT scanner system according to the presenttechnology, the vertical walls that define septa (and/or through-holesthereof) of 2D ASG 3410 may be aligned toward a focal spot of the x-raysource providing an x-ray emission path 3950 for use in imaging theobject 3940 positioned in the opening 3945. Such alignment may beprovided in system 3900 by the O-shaped gantry 3930 configuration havingthe aforementioned centered geometry, or alternatively or additionallyby the O-shaped gantry 3930 incorporating in the techniques of method3700 according to the present technology.

In some embodiments, the vertical walls defining septa (and/orthrough-holes thereof) of 2D ASG 3410 may be at least partiallyunaligned with the detector pixels of FPD 3920. For example, and withoutlimitation, the vertical walls of 2D ASG 3410 need not be parallel tothe detector pixel array of FPD 3920. However, in other embodiments, thevertical walls of 2D ASG 3410 can be aligned with the detector pixelarray of PFD 3920.

In some embodiments of system 3900, as well as of system(s) 3600 and/or3800 described above in Example 12, the 2D ASG 3410 may be fabricated asa single module. Alternatively, the 2D ASG 3410 may be fabricated assmaller modules that can be integrated together. In some embodiments,the 2D ASG 3410 may be placed directly on the x-ray absorbing sensorlayer of FPD 3920. In other embodiments, a gap may exist between 2D ASG3410 and the x-ray absorbing sensor layer of FPD 3920. As used in one ormore of system(s) 3600 and 3800 according to the present technology, the2D ASG 3410 may include any of the various design features disclosedherein to provide a number of advantageous technical effects toclinicians performing diagnostic x-ray imaging for the benefit of humanor veterinarian patients.

In some embodiments, the controller or computer 3990 of system 3900 maybe programmed, or otherwise configured, to perform or implement at leastin part any or all of the image processing methods and techniques asdescribed herein according to the present technology. FIG. 39G presentsa method 3902 for processing x-ray images that may be used with thesystem 3900 of FIGS. 39A-39F, according to some embodiments of thepresent technology.

Method 3902 provides an image data correction process that may beutilized with, for example and without limitations, one or more of thesystem(s) 3600, 3800 and 3900 according to the present technology.Method 3902 may, in some embodiments, be further employed with any ofthe various x-ray imaging devices and systems disclosed herein. A 2D ASG(e.g., 3410) may not stop all scattered x-rays. Some of the scatteredx-rays from an x-ray source (e.g., 3910) may be transmitted through the2D ASG. This issue is address by method 3902 according to the presenttechnology with a residual scatter correction process. Correction of theresidual scatter intensity according to the present technology may beaccomplished, at least in part, by measuring signal intensity variationsin the FPD or detector pixel array to facilitate an improvement in thequality of the x-ray image Method 3902 may minimize or otherwiseadvantageously reduces image artifacts in CT scanner images that can beintroduced by a 2D ASG (e.g., 3410) associated with an x-ray detector(e.g., FPD 3410).

In CT and CBCT systems, gantry flex can result a change in position ofthe x-ray source in relation to the x-ray detector while the gantry isrotating around the patient during a CT scan. Gantry flex can degradethe performance of x-ray imaging techniques. Use of an O-shaped gantry(e.g., 3930) can mitigate and suppress the gantry flex effect.Accordingly, method 3902 may be uniquely suited to use with x-rayimaging systems that include an O-shaped gantry (e.g., system 3900).

An O-shaped gantry is structurally more robust as compared to a C-shapedgantry, thereby minimizing the negative effects of x-ray imaging arisingfrom gantry flex. C-shaped gantry may be utilized with known portableand/or compact x-ray imaging devices and devices for the space-savingbenefits. Yet, O-shaped gantries tend to take up more space and may beheavier than systems using C-shaped gantries. The incorporation of theFPD 3920 and 2D ASG 3410 in the system 3900 according to the presenttechnology has hitherto not been considered in the art for at leastthese reasons, and as described in greater detail below, the results interms of imaging quality from a compact and/or portable x-ray imagingsystem according to the present technology are unexpected. As furtherdescribed below, the systems and methods according to the presenttechnology yield high quality diagnostic images that are comparable tostate of the art devices and systems that may not be portable and/orcompact.

In some embodiments, the O-shaped gantry 3930 of system 3900 may have aslip-ring design to transfer data and power without cables.Alternatively, O-shaped gantry 3930 may have cables (rather than slipring) to transfer data and power. In some embodiments, data transfer maybe accomplished wirelessly either instead of, or in addition to, thewired data connections. In an example of system 3900, a maximum rotationspeed of the O-shaped gantry 3930 will be 2-3 seconds per rotation, toreduce G-forces and hence, reduce structural weight and complexity. Bycomparison, in at least some known O-shaped CT gantries, rotationsspeeds approach 0.2 seconds per rotation, requiring a heavy andprecisely balanced gantry structure to handle high G forces. A lowerO-shaped gantry 3930 rotation speed in the system 3900 according to thepresent technology also allows protecting structural integrity of theFPD 3920 and the 2D ASG 3410. Furthermore, the ability to tilt theO-shaped 3930 in the CT scanner system according to the presenttechnology was hitherto challenging in at least some known portableand/or compact CT scanners to accommodate patient who are unable to layflat or nearly horizontal.

In some embodiments configured for imaging the head and extremities, aradial distance between the x-ray source 3910 and the FPD 3920 may befrom 60 cm to 100 cm. X-ray source 3910 to isocenter distance (SIsoD) insystem 3900 may be from 30 to 60 cm. The human brain is about 15-16 cmin length in the craniocaudal direction. To scan the whole brain in onerotation of O-shaped gantry 3910, a width of 25 to 30 cm for FPD 3920 inthe craniocaudal direction is needed. However, the use of such a largedetector also requires a large 2D ASG 3410, which can increase systemcomplexity. In some examples, use of a 20 cm wide FPD 3920 provides asufficient width to scan the whole brain in a 9-12 cm wide FOV in thecranio-caudal direction, and with high quality imaging results withoutundue complexity issues for system 3900.

In some embodiments, two scans may be performed using system 3900 toincrease the FOV, where the gantry is displaced about 10 cm in thecranio-caudal direction for the second scan, allowing a total coverageof 20 cm. To perform the second scan, gantry can be shiftedelectromechanically in the cranio-caudal direction, and multiple CTscans can be performed to increase craniocaudal coverage.

At least some known CBCT systems employ flat panel detectors with pulsedx-ray source and low frame acquisition rates, which can lead to longscan durations. To reduce scan duration, compact and/or portableembodiments of the CT scanner system 3900 according to the presenttechnology may employ continuous x-ray exposure at high frame rates. Toachieve high frame rates, pixel size may be increased to 0.4-0.6 mm, bybinning pixels. With this approach, frame rates can be increased to 100frame/second or more, allowing acquisition of 600-700 frames in 6-7seconds to in one rotation of the O-shaped gantry 3930. For perfusion CTscans, even higher frame rates may be needed to complete a scan in 3seconds or less. This can be achieved using the system 3900 according tothe present technology by reducing the readout area of the FPD 3920 inthe craniocaudal direction. The increase in frame rate will beproportional to the reduction in readout area.

In some embodiments, method 3902 may implement an image correctionalgorithm to correct residual scatter not stopped by the antiscattergrid for an image of the object 3940. Correction of the residual scatterintensity according to the present technology may be accomplished, atleast in part, by measuring signal intensity variations in the FPD ordetector pixel array to facilitate an improvement in the quality of thex-ray image. Referring now to FIG. 39G, method 3902 may include the stepof receiving 3904, from the FPD 3920, data representative of an image ofthe object 3940 to facilitate an x-ray image of at least a portion ofthe object to be generated based on the data representative of the imageof the object. In an example, method 3902 may also include the step ofgenerating 3906 the x-ray image of at least a portion of the object3940. Method 3902 may include the step of estimating 3908 a residualscatter intensity reaching the FPD 3920.

In some embodiments, method 3902 may include the step of correcting 3910the residual scatter intensity to facilitate an improvement in a qualityof the x-ray image of at least a portion of the object 3940. Correctionof the residual scatter intensity according to the present technologymay be accomplished, at least in part, by measuring signal intensityvariations in the FPD or detector pixel array to facilitate animprovement in the quality of the x-ray image. In one example, thecorrecting 3910 step of method 3902 may include correcting 3912 theresidual scatter intensity based on a pattern of signal intensityvariations in projections of images of the object acquired by thedetector pixel array.

In some embodiments, the estimating 3908 step of method 3902 may includedetermining 3914 a change in a ratio of an image intensity underneath afootprint of the 2D ASG 3410 to an image intensity in one or more of theopen-ended channels of 2D ASG 3410 in the absence of the object 3940positioned between the x-ray source 3910 and the FPD 3920. In someembodiments, the estimating 3908 step of method 3902 may includedetermining 3916 a change in the ratio of the image intensity underneaththe footprint of the 2D ASG 3410 to the image intensity in one or moreof the open-ended channels of 2D ASG 3410 in the presence of the object3940 positioned between the x-ray source 3910 and the FPD 3920. Method3902 may include either or both of the above-described determiningstep(s) 3914 and 3916.

In some embodiments, system 3900 may include one or more non-transitorycomputer readable media 3994 operably coupled to the controller orcomputer 3990. The medium or media may have method 3902-related programinstructions stored thereon. The program instructions may beprocessor-readable and executable code (e.g., software or firmware).When executed by, for example and without limitation, one or moreprocessors of the at least one controller or computer 3990, the programinstructions may cause the at least one of the above-describedprocesses, or steps, of method 3902 to be performed, at least in part,by the portable CT scanner system 3900, or similarly, system(s) 3600and/or 3800.

FIG. 39H presents CT scan images obtained using the CT scanner system3900 of FIGS. 39A-39F as compared to images obtained using known CTscanners. CT imaging experiments demonstrated an improvement in imagequality of the portable and/or compact system 3900 according to thepresent technology as compared to known CT scanners that utilize FPD.FIG. 39H includes three pairs of images from: panel (1) a state of theart FDA approved CT with FPD on the left hand side, (2) the CT scannersystem 3900 according to the present technology in the middle, and (3) a“gold standard” helical CT scanner—the Philips® Healthcare 16 slice CTscanner.

In the images of panel (1) of FIG. 39H, image artifacts are visible.Such image quality is unacceptable in the context of soft tissueimaging. In the images of panel (2) of FIG. 39H, shading artifacts aresuppressed, and low contrast objects are clearly visible. No datacorrection methods were applied besides scatter correction to the panel(2) images. As can be seen by comparing the panel (2) images to thepanel (3) images of FIG. 39H, the images generated using system 3900according to the present technology are comparable in quality to the“gold standard” Philips® Healthcare 16 slice CT scanner.

Conventional CT detectors (and associated antiscatter grids) may requiremany CT gantry rotations to scan the anatomy of interest. Making manygantry rotations has many adverse effects. The gantry has to rotate veryfast, and the gantry has to be built strong enough to withstand theg-forces. This approach not only increases the costs of the system, italso makes the system heavier. Requiring many gantry rotations may alsorequires a heavier duty x-ray tube and power generator to be able togenerate x-rays during many gantry rotations. Furthermore, conventionalCT detectors are custom designed for each system, which increase thecosts.

In the x-ray imaging systems according to the present technology, use ofa flat panel detector with a 2D antiscatter grid (and associatedmethods), to make the portable CT scanner lighter, more portable, andcheaper.

The use of antiscatter grids in a conventional CT scanner is a knownapproach. However, in the systems according to the present technology, asubstantially different detector (i.e. flat panel detector) is used inthe compact and/or portable CT scanner. The use of flat panel detectorwith 2D ASG brings substantially different challenges.

A known portable and conventional CT detector for brain imaging canimage a 1 cm thick slice of brain per gantry rotation. Whereas thesystems according to the present technology (e.g., CT scanner system3900) uses an FPD and 2D ASG, along with methods for successfulintegration of the FPD and the 2D ASG, to image a 10+ cm thick slice ofbrain in one gantry rotation. Given that the brain is 15-16 cm long inthe cranio-caudal direction, the known CT detector needs to complete 10+gantry rotations to scan the whole brain. As a result, the known CTscanner requires: (a) fast rotating gantry (1-sec per rotation), whichmakes the gantry heavy and expensive; (b) a mechanism to move thescanner on the room floor, to scan the whole brain (without thismechanism, the known CT scanner can scan only 1 cm thick section of thebrain; and (c) high power x-ray generator and x-ray tube to providex-ray generation sufficient for 15-16 gantry rotations. Some knownexamples do not include the aforementioned mechanism to more the scanneron the room floor, however. The devices, systems and methods accordingto the present technology aim to address the aforementioned issues withthe known CT scanner, as in the present technology (e.g., system 3900)does not require a fast-rotating gantry, it scan the whole brain withone or two gantry rotations, and therefore it does not need to move onthe room floor, and it can operate using less x-ray power.

Another known x-ray imaging system is portable, uses a flat paneldetector, and can image the whole head with one gantry rotary rotation.However, this known system does not use a 2D antiscatter grid to addressthe scatter problem, as described herein. This known system thereforedoes not provide sufficient image quality for visualizing brain, and ituses a C-shaped gantry which is not optimal for implementation of a 2Dantiscatter grid due to, for example, gantry flex. When a C-shapedgantry rotates, its arms flex slightly due to gravity. Such a flexing inthe arms, cause image artifacts when 2D antiscatter grid is in place.The present technology (e.g., system 3900) may use an O-shaped gantry,which reduces the gantry flex significantly and reduces image artifactswhen 2D antiscatter grid is used. Moreover, the use of an O-shapedgantry (hence the reduced gantry flex) is better suited for the datacorrection methods according to the present technology (e.g., method3902) needed for 2D antiscatter grid implementation. A C-shaped gantryrotates relatively slow (˜8-20 secs per 360 degree rotation), whereasthe O-shaped gantry 3930 of system 3900 may rotate as fast as 3 secondsper rotation. Operating a C-shaped gantry at higher rotation speeds maypresent safety concerns. If a C-shaped gantry hits the operator or thepatient (e.g., due to unintentional movement of the patient), it cancause serious injury. These concerns are eliminated, or at leastmitigated, using the O-shaped gantry 3930 in system 3900, for example.

Example 14

A commercially available CBCT scanner was modified to incorporate the 2Dantiscatter grid hardware according to the present technology and wasused to acquire images. FIGS. 40A and 40B present images in comparisonbetween an ultracompact CT (uCT) scanner with a 2D antiscatter gridaccording to some embodiments of the present technology (e.g., system3900 of FIGS. 39A-39F) and a helical CT scanner. The latter system isreferred to in FIG. 40B as a “gold standard.” The images in FIG. 40Awere acquired using the uCT scanner with the 2D antiscatter gridaccording to the present technology (e.g., 2D ASG 3410), and the imagesin FIG. 40B were acquired by using the gold standard helical CT scanner.Comparing the pair of images of two different slices of a human head inFIG. 40A to the pair of images of the same physiological features inFIG. 40B, it can be seen that the uCT scanner using the 2D antiscattergrid demonstrates noninferiority of the present technology with respectto the gold standard approach.

Example 15

FIGS. 41A and 41B present a comparison of images obtained using a CBCTsystem and processed without and with, respectively, the techniques ofthe method 3902 shown in FIG. 39G. As can be seen in the pair of imagesin FIG. 41A, sever shading artifacts are apparent. Such artifacts are anindication of poor scatter suppression and poor image quality. Movingnow to the pair of images in FIG. 41B, application of the correctionalgorithm by method 3902 according to the present technology results insubstantially fewer artifacts, which indicates a higher image quality.

Example 16

FIGS. 42A and 42B present a comparison of images obtained using the goldstandard CT and a CBCT scanner (e.g., system 3900), respectively,according to the present technology. The pair of images shown in FIG.42B, acquired with the 2D antiscatter grid according to the presenttechnology (e.g., 2D ASG 3410) are expected to be similar to the pair ofimages of FIG. 42A using the gold standard system. Qualitatively, theseimages indicate that the approach according to the present technology(e.g., system 3900) provides comparable image quality when compared togold standard CT scans.

Example 17

FIGS. 43A, 43B and 43C present simulated images in comparison todemonstrate the effect of CT gantry shape on gantry flex and 2Dantiscatter grid performance. The insets in each of FIGS. 43A-43C arethe zoomed-in central portion of each CT image. The simulated CT imageof FIG. 43A was produced using an O-shaped gantry (e.g., 3930) withassociated 2D ASG (e.g., 3410). The simulated CT image of FIGS. 43B and43C were produced using a C-shaped gantry (3930) with associated 2D ASG(e.g., 3410).

The image of FIG. 43A demonstrates image quality when there is no gantryflex as in an O-shaped gantry. The simulated CT images of FIGS. 43B and43C with small, and medium, amounts of gantry flex, respectively, in theC-shaped gantry. An O-shaped gantry (e.g., 3930) minimizes the gantryflex when the gantry is rotating around the patient. As a result, a 2Dantiscatter grid's (e.g., 2D ASG 3410) shadows can be corrected using aconventional flat field correction technique, and ring artifacts in CTimages are avoided. Whereas in a C-shaped gantry, the grid shadows arenot properly suppressed, causing ring artifacts in CT images. Dependingon the magnitude of gantry flex, the intensity of ring artifacts willalso vary. Notably, x-ray imaging systems that do not utilize 2D ASDsmay experience less of an issue arising from gantry flex. Accordingly,the technically advantageous features of the systems and methods (e.g.,system 3900 and method 3902) according to the present technology canprevent, or at least mitigate, the aforementioned negative effects ofgantry flex that were hitherto extant in at least some known x-rayimaging systems.

FIG. 44 presents a diagrammatic representation of a machine, in theexample form, of a computer system 700 within which a set ofinstructions, for causing the machine to implement or otherwise performany one or more of the techniques and methodologies of the presenttechnology described herein, may be executed. In the example of FIG. 44, the computer system 700 includes a processor, memory, non-volatilememory, and an interface device. Various common components (e.g., cachememory) are omitted for illustrative simplicity. The computer system 700is intended to illustrate a hardware device on which any of thecomponents depicted in the foregoing examples and embodiments of thepresent technology can be implemented. For example, the computer system700 can be, or may include, an image processing system or a motorcontrol system. The computer system 700 can be of any applicable knownor convenient type. The components of the computer system 700 can becoupled together via a bus or through some other known or convenientdevice.

The processor of computer system 700 may be, for example, a conventionalmicroprocessor such as an INTEL PENTIUM microprocessor or MOTOROLA POWERPC microprocessor. One of skill in the relevant art will recognize thatthe terms “machine-readable (storage) medium” or “computer-readable(storage) medium” include any type of device that is accessible by theprocessor. In some embodiment, these storage media are embodied innon-transitory computer-readable media that can store programinstructions (e.g., as software or firmware) which, when executed by oneor more processors of the disclosed technology, cause the processor(s)or other controller means to implement, execute, or otherwise facilitateperformance of the various algorithms and methods disclosed herein.

In computer system 700, the memory is coupled to the processor by, forexample, a bus. The memory can include, by way of example but notlimitation, random access memory (RAM), such as dynamic RAM (DRAM) andstatic RAM (SRAM). The memory can be local, remote, or distributed.

The bus of computer system 700 also couples the processor to thenon-volatile memory and drive unit. The non-volatile memory is often amagnetic floppy or hard disk, a magnetic-optical disk, an optical disk,a read-only memory (ROM), such as a CD-ROM, EPROM, or EEPROM, a magneticor optical card, or another form of storage for large amounts of data.Some of this data is often written, by a direct memory access process,into memory during execution of software in the computer system 700. Thenon-volatile storage can be local, remote, or distributed. Thenon-volatile memory is optional because systems can be created with allapplicable data available in memory. An embodiment of computer system700 will usually include at least a processor, memory, and a device(e.g., a bus) coupling the memory to the processor.

Software or firmware utilized by computer system 700 may be stored inthe non-volatile memory and/or the drive unit. Indeed, for largeprograms, it may not even be possible to store the entire program in thememory. Nevertheless, it should be understood that for software and/orfirmware to run, if necessary, it is moved to a computer readablelocation appropriate for processing, and for illustrative purposes, thatlocation is referred to as the memory in this paper. Even when softwareis moved to the memory for execution, the processor will typically makeuse of hardware registers to store values associated with the software,and local cache that, ideally, serves to speed up execution. As usedherein, firmware or a software program is assumed to be stored at anyknown or convenient location (from non-volatile storage to hardwareregisters) when the software program is referred to as “implemented in acomputer-readable medium”. A processor is considered to be “configuredto execute a program” when at least one value associated with theprogram is stored in a register readable by the processor.

The bus also couples the processor to the network interface device ofcomputer system 700. The interface can include one or more of a modem ornetwork interface. It will be appreciated that a modem or networkinterface can be considered to be part of the computer system. Theinterface can include an analog modem, ISDN modem, cable modem, tokenring interface, satellite transmission interface (e.g., “direct PC”), orother interfaces for coupling a computer system (e.g., 700) to othercomputer systems. The interface can include one or more input and/oroutput (I/O) devices. The I/O devices can include, by way of example butnot limitation, a keyboard, a mouse or other pointing device, diskdrives, printers, a scanner, and other input and/or output devices,including a display device. The display device can include, by way ofexample but not limitation, a cathode ray tube (CRT), liquid crystaldisplay (LCD), or some other applicable known or convenient displaydevice. For simplicity, it is assumed that controllers of any devicesnot depicted in the example of FIG. 44 reside in the interface.

In operation, the computer system 700 can be controlled by operatingsystem software that includes a file management system, such as a diskoperating system. One example of operating system software withassociated file management system software is the family of operatingsystems known as WINDOWS from MICROSOFT Corporation of Redmond, Wash.,and their associated file management systems. Another example ofoperating system software with its associated file management systemsoftware is the LINUX operating system and its associated filemanagement system. The file management system is typically stored in thenon-volatile memory and/or drive unit and causes the processor toexecute the various acts required by the operating system to input andoutput data and to store data in the memory, including storing files onthe non-volatile memory and/or drive unit.

Some portions of the detailed description may be presented in terms ofalgorithms and symbolic representations of operations on data bitswithin a computer memory. These algorithmic descriptions andrepresentations are the means used by those skilled in the dataprocessing arts to most effectively convey the substance of their workto others skilled in the art. An algorithm is here, and generally,conceived to be a self-consistent sequence of operations leading to adesired result. The operations are those requiring physicalmanipulations of physical quantities. Usually, though not necessarily,these quantities take the form of electrical or magnetic signals capableof being stored, transferred, combined, compared, and otherwisemanipulated. It has proven convenient at times, principally for reasonsof common usage, to refer to these signals as bits, values, elements,symbols, characters, terms, numbers, or the like.

It should be borne in mind, however, that all of these and similar termsare to be associated with the appropriate physical quantities and aremerely convenient labels applied to these quantities. Unlessspecifically stated otherwise, as apparent from the followingdiscussion, it is appreciated that throughout the description,discussions utilizing terms such as “processing” or “computing” or“calculating” or “determining” or “displaying” or the like, refer to theaction and processes of a computer system, or similar electroniccomputing device, that manipulates and transforms data represented asphysical (electronic) quantities within the computer system's registersand memories into other data similarly represented as physicalquantities within the computer system memories or registers or othersuch information storage, transmission or display devices.

The algorithms and displays presented herein are not inherently relatedto any particular computer or other apparatus. Various general-purposesystems may be used with programs in accordance with the teachingsherein, or it may prove convenient to construct more specializedapparatus to perform the methods of some embodiments. The requiredstructure for a variety of these systems will appear from thedescription below. In addition, the techniques are not described withreference to any particular programming language, and variousembodiments may thus be implemented using a variety of programminglanguages.

In alternative embodiments, the machine operates as a standalone deviceor may be connected (e.g., networked) to other machines. In a networkeddeployment, the machine may operate in the capacity of a server or aclient machine in a client-server network environment or as a peermachine in a peer-to-peer (or distributed) network environment.

The machine may be a server computer, a client computer, a personalcomputer (PC), a tablet PC, a laptop computer, a set-top box (STB), apersonal digital assistant (PDA), a cellular telephone, an IPHONE, aBLACKBERRY, a processor, a telephone, a web appliance, a network router,switch or bridge, or any machine capable of executing a set ofinstructions (sequential or otherwise) that specify actions to be takenby that machine.

While the machine-readable medium or machine-readable storage medium isshown in an exemplary embodiment to be a single medium, the term“machine-readable medium” and “machine-readable storage medium” shouldbe taken to include a single medium or multiple media (e.g., acentralized or distributed database, and/or associated caches andservers) that store the one or more sets of instructions. The term“machine-readable medium” and “machine-readable storage medium” shallalso be taken to include any medium that is capable of storing, encodingor carrying a set of instructions for execution by the machine and thatcause the machine to perform any one or more of the methodologies of thepresently disclosed technique and innovation.

In general, the routines executed to implement the embodiments of thedisclosure, may be implemented as part of an operating system or aspecific application, component, program, object, module or sequence ofinstructions referred to as “computer programs.” The computer programstypically comprise one or more instructions set at various times invarious memory and storage devices in a computer, and that, when readand executed by one or more processing units or processors in acomputer, cause the computer to perform operations to execute elementsinvolving the various aspects of the disclosure.

Moreover, while embodiments have been described in the context of fullyfunctioning computers and computer systems, those skilled in the artwill appreciate that the various embodiments are capable of beingdistributed as a program product in a variety of forms, and that thedisclosure applies equally regardless of the particular type of machineor computer-readable media used to actually effect the distribution.

Further examples of machine-readable storage media, machine-readablemedia, or computer-readable (storage) media include but are not limitedto recordable type media such as volatile and non-volatile memorydevices, floppy and other removable disks, hard disk drives, opticaldisks (e.g., Compact Disk Read-Only Memory (CD ROMS), Digital VersatileDisks, (DVDs), etc.), among others, and transmission type media such asdigital and analog communication links.

Unless the context clearly requires otherwise, throughout thedescription and the claims, the words “comprise,” “comprising,” and thelike are to be construed in an inclusive sense, as opposed to anexclusive or exhaustive sense; that is to say, in the sense of“including, but not limited to.” As used herein, the terms “connected,”“coupled,” or any variant thereof, means any connection or coupling,either direct or indirect, between two or more elements; the coupling ofconnection between the elements can be physical, logical, or acombination thereof. Additionally, the words “herein,” “above,” “below,”and words of similar import, when used in this application, shall referto this application as a whole and not to any particular portions ofthis application. Where the context permits, words in the above detaileddescription using the singular or plural number may also include theplural or singular number, respectively. The word “or,” in reference toa list of two or more items, covers all of the following interpretationsof the word: any of the items in the list, all of the items in the list,and any combination of the items in the list.

The above detailed description of embodiments of the disclosure is notintended to be exhaustive or to limit the teachings to the precise formdisclosed above. While specific embodiments of, and examples for, thedisclosure are described above for illustrative purposes, variousequivalent modifications are possible within the scope of thedisclosure, as those skilled in the relevant art will recognize. Forexample, while processes or blocks are presented in a given order,alternative embodiments may perform routines having steps, or employsystems having blocks, in a different order, and some processes orblocks may be deleted, moved, added, subdivided, combined, and/ormodified to provide alternative or subcombinations. Each of theseprocesses or blocks may be implemented in a variety of different ways.Also, while processes or blocks are, at times, shown as being performedin a series, these processes or blocks may instead be performed inparallel, or may be performed at different times. Further, any specificnumbers noted herein are only examples: alternative implementations mayemploy differing values or ranges.

The teachings of the disclosure provided herein can be applied to othersystems, not necessarily the system described above. Likewise, theelements and acts of the various embodiments described above can becombined to provide further embodiments.

These and other changes can be made to the disclosure in light of theabove detailed description. While the above description describescertain embodiments of the disclosure, and describes the best modecontemplated, no matter how detailed the above appears in text, theteachings can be practiced in many ways. Details of the system may varyconsiderably in its implementation details, while still beingencompassed by the subject matter disclosed herein. As noted above,particular terminology used when describing certain features or aspectsof the disclosure should not be taken to imply that the terminology isbeing redefined herein to be restricted to any specific characteristics,features, or aspects of the disclosure with which that terminology isassociated. In general, the terms used in the following claims shouldnot be construed to limit the disclosure to the specific embodimentsdisclosed in the specification, unless the above detailed descriptionsection explicitly defines such terms. Accordingly, the actual scope ofthe disclosure encompasses not only the disclosed embodiments, but alsoall equivalent ways of practicing or implementing the disclosure underthe claims.

Any patents or patent applications and other references noted above,including any that may be listed in accompanying filing papers, areincorporated herein by reference in their entireties. Aspects of thedisclosure can be modified, if necessary, to employ the systems,functions, and concepts of the various references described above toprovide yet further embodiments of the disclosure.

While certain aspects of the disclosure are presented below in certainclaim forms, the inventors contemplate the various aspects of thedisclosure in any number of claim forms. For example, while only oneaspect of the disclosure is recited as a means-plus-function claim under35 U.S.C. § 112(f), other aspects may likewise be embodied as ameans-plus-function claim, or in other forms, such as being embodied ina computer-readable medium. (Any claims intended to be treated under 35U.S.C. § 112(f) will begin with the words “means for”.) Accordingly, theapplicant reserves the right to add additional claims after filing theapplication to pursue such additional claim forms for other aspects ofthe disclosure.

The detailed description provided herein may be applied to otherdevices, systems and methods, not necessarily only the devices, systemsand methods described above. The elements and acts of the variousexamples and embodiments described above can be combined to providefurther implementations of the invention. Some alternativeimplementations of the invention may include not only additionalelements to those implementations noted above, but also may includefewer elements. These and other changes can be made to the invention inlight of the above detailed description. While the above descriptiondefines certain examples of the invention, and describes the best modecontemplated, no matter how detailed the above appears in text, theinvention can be practiced in many ways. Details of the system may varyconsiderably in its specific implementation, while still beingencompassed by the invention disclosed herein. As noted above,particular terminology used when describing certain features or aspectsof the invention should not be taken to imply that the terminology isbeing redefined herein to be restricted to any specific characteristics,features, or aspects of the invention with which that terminology isassociated. In general, the terms used in the following claims shouldnot be construed to limit the invention to the specific examplesdisclosed in the specification, unless the above detailed descriptionsection explicitly defines such terms. Accordingly, the actual scope ofthe invention encompasses not only the disclosed examples, but also allequivalent ways of practicing or implementing the invention.

The illustrations of the embodiments described herein are intended toprovide a general understanding of the structure of the variousembodiments. The illustrations are not intended to serve as a completedescription of all of the elements and features of apparatus and systemsthat utilize the structures or methods described herein. Many otherembodiments may be apparent to those of skill in the art upon reviewingthe disclosure. Other embodiments may be utilized and derived from thedisclosure, such that structural and logical substitutions and changesmay be made without departing from the scope of the disclosure.Moreover, although specific embodiments have been illustrated anddescribed herein, it should be appreciated that any subsequentarrangement designed to achieve the same or similar purpose may besubstituted for the specific embodiments shown.

This disclosure is intended to cover any and all subsequent adaptationsor variations of various embodiments. Combinations of the aboveembodiments can be made, and other embodiments not specificallydescribed herein will be apparent to those of skill in the art uponreviewing the description. Additionally, the illustrations are merelyrepresentational and may not be drawn to scale. Certain proportionswithin the illustrations may be exaggerated, while other proportions maybe reduced. Accordingly, the disclosure and the figures are to beregarded as illustrative and not restrictive.

REFERENCES

-   1. Altunbas, C. et al. (2014) “Su-D-12a-04: Investigation of a 2D    Antiscatter Grid for Flat Panel Detectors,” Med. Phys. 41(6),    124-124.-   2. Fetterly, K. A. and Schueler, B. A. (2007) “Experimental    Evaluation of Fiber-Interspaced Antiscatter Grids for Large Patient    Imaging with Digital X-Ray Systems,” Phys. Med. Biol. 52(16),    4863-4880.-   3. Fetterly, K. A. and Schueler, B. A. (2009) “Physical Evaluation    of Prototype High-Performance Anti-Scatter Grids: Potential for    Improved Digital Radiographic Image Quality,” Phys. Med. Biol.    54(2), N37-42.-   4. Siewerdsen, J. H. and Jaffray, D. A. (2001) “Cone-Beam Computed    Tomography with a Flat-Panel Imager: Magnitude and Effects of X-Ray    Scatter,” Med. Phys. 28(2), 220-231.-   5. Ruhrnschopf, E.-P. and Klingenbeck, K. (2011) “A General    Framework and Review of Scatter Correction Methods in X-Ray    Cone-Beam Computerized Tomography. Part 1: Scatter Compensation    Approaches,” Med. Phys. 38(7), 4296-4311.-   6. Angle, J. F. (2013) “Cone-Beam Ct: Vascular Applications,” Tech.    Vasc. Interv. Radiol. 16(3), 144-149.-   7. De Vos, W. et al. (2009) “Cone-Beam Computerized Tomography    (Cbct) Imaging of the Oral and Maxillofacial Region: A Systematic    Review of the Literature,” Int. J. Oral Maxillofac. Surg. 38(6),    609-625.-   8. Van de Kelft, E. et al. (2012) “A Prospective Multicenter    Registry on the Accuracy of Pedicle Screw Placement in the Thoracic,    Lumbar, and Sacral Levels with the Use of the O-Arm Imaging System    and Stealthstation Navigation,” Spine 37(25), E1580-E1587.-   9. Simpson, D. R. et al. (2010) “A Survey of Image-Guided Radiation    Therapy Use in the United States,” Cancer 116(16), 3953-3960.-   10. Jaffray, D. A. (2012) “Image-Guided Radiotherapy: From Current    Concept to Future Perspectives,” Nature Reviews Clinical Oncology    9(12), 688-699.-   11. Yang, H. et al. (2013) “Replanning During Intensity Modulated    Radiation Therapy Improved Quality of Life in Patients with    Nasopharyngeal Carcinoma,” International Journal of Radiation    Oncology* Biology* Physics 85(1), e47-e54.-   12. Castadot, P. et al. (2010) “Adaptive Radiotherapy of Head and    Neck Cancer,” Semin. Radiat. Oncol. 20(2), 84-93.-   13. Schwartz, D. L. et al. (2013) “Adaptive Radiotherapy for Head    and Neck Cancer—Dosimetric Results from a Prospective Clinical    Trial,” Radiother. Oncol. 106(1), 80-84.-   14. Castadot, P. et al. (2011) “Adaptive Functional Image-Guided    Imrt in Pharyngo-Laryngeal Squamous Cell Carcinoma: Is the Gain in    Dose Distribution Worth the Effort?,” Radiother. Oncol. 101(3),    343-350.-   15. Hansen, E. K. et al. (2006) “Repeat Ct Imaging and Replanning    During the Course of Imrt for Head-and-Neck Cancer,” International    Journal of Radiation Oncology* Biology* Physics 64(2), 355-362.-   16. Barker, J. L. et al. (2004) “Quantification of Volumetric and    Geometric Changes Occurring During Fractionated Radiotherapy for    Head-and-Neck Cancer Using an Integrated Ct/Linear Accelerator    System,” International Journal of Radiation Oncology* Biology*    Physics 59(4), 960-970.-   17. Foroudi, F. et al. (2011) “Online Adaptive Radiotherapy for    Muscle-Invasive Bladder Cancer: Results of a Pilot Study,”    International Journal of Radiation Oncology* Biology* Physics 81(3),    765-771.-   18. Tyagi, N. et al. (2011) “Daily Online Cone Beam Computed    Tomography to Assess Interfractional Motion in Patients with Intact    Cervical Cancer,” International Journal of Radiation Oncology*    Biology* Physics 80(1), 273-280.-   19. Bertelsen, A. et al. (2011) “Radiation Dose Response of Normal    Lung Assessed by Cone Beam Ct—a Potential Tool for Biologically    Adaptive Radiation Therapy,” Radiother. Oncol. 100(3), 351-355.-   20. Kwint, M. et al. (2014) “Intra Thoracic Anatomical Changes in    Lung Cancer Patients During the Course of Radiotherapy,” Radiother.    Oncol. 113(3), 392-397.-   21. Weiss, E. et al. (2010) “Clinical Evaluation of Soft Tissue    Organ Boundary Visualization on Cone-Beam Computed Tomographic    Imaging,” IJROBP 78(3), 929-936.-   22. Lütgendorf-Caucig, C. et al. (2011) “Feasibility of Cbct-Based    Target and Normal Structure Delineation in Prostate Cancer    Radiotherapy: Multi-Observer and Image Multi-Modality Study,”    Radiother. Oncol. 98(2), 154-161.-   23. Hou, J. et al. (2011) “Deformable Planning Ct to Cone-Beam Ct    Image Registration in Head-and-Neck Cancer,” Med. Phys. 38(4),    2088-2094.-   24. Veiga, C. et al. (2014) “Toward Adaptive Radiotherapy for Head    and Neck Patients:-   Feasibility Study on Using Ct-to-Cbct Deformable Registration for    “Dose of the Day” Calculations,” Med. Phys. 41(3), 031703.-   25. Møller, D. S. et al. (2014) “Adaptive Radiotherapy of Lung    Cancer Patients with Pleural Effusion or Atelectasis,” Radiother.    Oncol. 110(3), 517-522.-   26. Chen, A. M. et al. (2014) “Clinical Outcomes among Patients with    Head and Neck Cancer Treated by Intensity—Modulated Radiotherapy    with and without Adaptive Replanning,” Head Neck 36(11), 1541-1546.-   27. Kyriakou, Y. and Kalender, W. (2007) “Efficiency of Antiscatter    Grids for Flat-Detector Ct,” Phys. Med. Biol. 52(20), 6275-6293.-   28. Lazos, D. and Williamson, J. F. (2010) “Monte Carlo Evaluation    of Scatter Mitigation Strategies in Cone-Beam Ct,” Med. Phys.    37(10), 5456-5470.-   29. Altunbas, C. (2014) “Image Corrections for Scattered Radiation,”    in Cone Beam Computed Tomography (Shaw, C. C., Ed.), pp 129-147, CRC    Press, Boca Raton, Fla.-   30. Ding, G. X. et al. (2007) “A Study on Adaptive Imrt Treatment    Planning Using Kv Cone-Beam Ct,” Radiother. Oncol. 85(1), 116-125.-   31. Guan, H. and Dong, H. (2009) “Dose Calculation Accuracy Using    Cone-Beam Ct (Cbct) for Pelvic Adaptive Radiotherapy,” Phys. Med.    Biol. 54(20), 6239.-   32. Fotina, I. et al. (2012) “Feasibility of Cbct-Based Dose    Calculation: Comparative Analysis of Hu Adjustment Techniques,”    Radiother. Oncol. 104(2), 249-256.-   33. Niu, T. et al. (2012) “Quantitative Cone-Beam Ct Imaging in    Radiation Therapy Using Planning Ct as a Prior: First Patient    Studies,” Med. Phys. 39(4), 1991-2000.-   34. Yang, Y. et al. (2007) “Evaluation of on-Board Kv Cone Beam Ct    (Cbct)-Based Dose Calculation,” Phys. Med. Biol. 52(3), 685.-   35. Ruhrnschopf, E. P. and Klingenbeck, K. (2011) “A General    Framework and Review of Scatter Correction Methods in X-Ray    Cone-Beam Computerized Tomography. Part 1: Scatter Compensation    Approaches,” Med. Phys. 38(7), 4296-4311.-   36. Ruhrnschopf, E.-P. and Klingenbeck, a. K. (2011) “A General    Framework and Review of Scatter Correction Methods in Cone Beam Ct.    Part 2: Scatter Estimation Approaches,” Med. Phys. 38(9), 5186-5199.-   37. Altunbas, M. C. et al. (2007) “A Post-Reconstruction Method to    Correct Cupping Artifacts in Cone Beam Breast Computed Tomography,”    Med. Phys. 34(7), 3109-3118.-   38. Liu, X. et al. (2006) “An Accurate Scatter Measurement and    Correction Technique for Cone Beam Breast Ct Imaging Using Scanning    Sampled Measurement (Ssm)Technique,” Proc. SPIE-Int. Soc. Opt. Eng.    6142, 614234.-   39. Siewerdsen, J. H. et al. (2004) “The Influence of Antiscatter    Grids on Soft-Tissue Detectability in Cone-Beam Computed Tomography    with Flat-Panel Detectors,” Med. Phys. 31(12), 3506-3520.-   40. Schafer, S. et al. (2012) “Antiscatter Grids in Mobile C-Arm    Cone-Beam Ct: Effect on Image Quality and Dose,” Med. Phys. 39(1),    153-159.-   41. Sisniega, A. et al. (2013) “Monte Carlo Study of the Effects of    System Geometry and Antiscatter Grids on Cone-Beam Ct Scatter    Distributions,” Med. Phys. 40(5), -.-   42. Wiegert, J. et al. (2004) “Performance of Standard Fluoroscopy    Antiscatter Grids in Flat-Detector-Based Cone-Beam Ct,” Proc.    SPIE-Int. Soc. Opt. Eng. 5368, 67-78.-   43. Lazos, D. et al. (2007) “Evaluation of Scatter Mitigation    Strategies for X-Ray Cone-Beam Ct: Impact of Scatter Subtraction and    Anti-Scatter Grids on Contrast-to-Noise Ratio,” (Jiang, H. and    Michael, J. F., Eds.), p 65101V, SPIE.-   44. Neitzel, U. (1992) “Grids or Air Gaps for Scatter Reduction in    Digital Radiography: A Model Calculation,” Med. Phys. 19(2),    475-481.-   45. Bonenkamp, J. G. and Boldingh, W. H. (1959) “Quality and Choice    of Potter Bucky Grids,” Acta Radiologica [Old Series] 52(3),    241-253.-   46. Santos, E. C. et al. (2006) “Rapid Manufacturing of Metal    Components by Laser Forming,” International Journal of Machine Tools    and Manufacture 46(12-13), 1459-1468.-   47. Gray, J. and Princehorn, J. “Htc Grids Improve Mammography    Contrast (White Paper),” (Inc., H., Ed.).-   48. Vogtmeier, G. et al. (2008) “Two-Dimensional Anti-Scatter Grids    for Computed Tomography Detectors,” (Jiang, H. and Ehsan, S., Eds.),    p 691359, SPIE.-   49. Kawrakow, I. and Rogers, D. (2000) “The Egsnrc Code System:    Monte Carlo Simulation of Electron and Photon Transport.”-   50. Yi, Y. et al. (2011) “Radiation Doses in Cone-Beam Breast    Computed Tomography: A Monte Carlo Simulation Study,” Med. Phys.    38(2), 589-597.-   51. Varian. (2010) “The Truebeam Technical Reference Guide-Volume 2:    Imaging,” Varian Medical Systems, Inc, Palo Alto, Calif.-   52. Johns, P. C. and Yaffe, M. (1982) “Scattered Radiation in Fan    Beam Imaging Systems,” Med. Phys. 9(2), 231-239.-   53. Abella, M. et al. (2012) “Software Architecture for Multi-Bed    Fdk-Based Reconstruction in X-Ray Ct Scanners,” Comput. Methods    Programs Biomed. 107(2), 218-232.-   54. Hsieh, J. (2009) in Computed Tomography: Principles, Design,    Artifacts, and Recent Advances, SPIE Press, Bellingham, Wash.-   55. Starman, J. et al. (2011) “Investigation into the Optimal Linear    Time-Invariant Lag Correction for Radar Artifact Removal,” Med.    Phys. 38(5), 2398-2411.-   56. Altunbas, C. et al. (2014) “Reduction of Ring Artifacts in Cbct:    Detection and Correction of Pixel Gain Variations in Flat Panel    Detectors,” Med. Phys. 41(9), 091913.-   57. Sharpe, M. B. et al. (2006) “The Stability of Mechanical    Calibration for a Kv Cone Beam Computed Tomography System Integrated    with Linear Acceleratora),” Med. Phys. 33(1), 136-144.-   58. Bissonnette, J. P. et al. (2008) “Quality Assurance for the    Geometric Accuracy of Cone-Beam Ct Guidance in Radiation Therapy,”    Int. J. Radiat. Oncol. Biol. Phys. 71(1 Suppl), S57-61.-   59. Gao, S. et al. (2014) “Evaluation of Isocal Geometric    Calibration System for Varian Linacs Equipped with on-Board Imager    and Electronic Portal Imaging Device Imaging Systems,” J. Appl.    Clin. Med. Phys. 15(3), 4688.-   60. Zheng, D. et al. (2011) “Bow-Tie Wobble Artifact: Effect of    Source Assembly Motion on Cone-Beam Ct,” Med. Phys. 38(5),    2508-2514.-   61. Jaffray, D. A. et al. (2002) “Flat-Panel Cone-Beam Computed    Tomography for Image-Guided Radiation Therapy,” International    Journal of Radiation Oncology, Biology, Physics 53(5), 1337-1349.-   62. Stankovic, U. et al. (2014) “Improved Image Quality of Cone Beam    Ct Scans for Radiotherapy Image Guidance Using Fiber-Interspaced    Antiscatter Grid,” Med. Phys. 41(6).-   63. Sijbers, J. and Andrei, P. (2004) “Reduction of Ring Artefacts    in High Resolution Micro-Ct Reconstructions,” Phys. Med. Biol.    49(14), N247.-   64. Star-Lack, J. et al. (2006) “Su-Ff-I-04: A Fast    Variable-Intensity Ring Suppression Algorithm,” Med. Phys. 33(6),    1997.-   65. Hatton, J. et al. (2009) “Cone Beam Computerized Tomography: The    Effect of Calibration of the Hounsfield Unit Number to Electron    Density on Dose Calculation Accuracy for Adaptive Radiation    Therapy,” Phys. Med. Biol. 54(15), N329.-   66. Altunbas, C. et al. (2013) “Dosimetric Errors During Treatment    of Centrally Located Lung Tumors with Stereotactic Body Radiation    Therapy: Monte Carlo Evaluation of Tissue Inhomogeneity    Corrections,” Med. Dosim. 38(4), 436-441.-   67. Miften, M. et al. (2001) “Comparison of Rtp Dose Distributions    in Heterogeneous Phantoms with the Beam Monte Carlo Simulation    System,” J. Appl. Clin. Med. Phys. 2(1), 21-31.-   68. Gayou, O. et al. (2007) “Patient Dose and Image Quality from    Mega-Voltage Cone Beam Computed Tomography Imaging,” Med. Phys.    34(2), 499-506.-   69. Fogliata, A. et al. (2007) “On the Dosimetric Behaviour of    Photon Dose Calculation Algorithms in the Presence of Simple    Geometric Heterogeneities: Comparison with Monte Carlo    Calculations,” Phys. Med. Biol. 52(5), 1363.-   70. Knöös, T. et al. (2006) “Comparison of Dose Calculation    Algorithms for Treatment Planning in External Photon Beam Therapy    for Clinical Situations,” Phys. Med. Biol. 51(22), 5785.-   71. Altunbas, C. et al. (2016) “We-Ab-207a-10: Transmission    Characteristics of a Two Dimensional Antiscatter Grid Prototype for    Cbct,” Med. Phys. 43(6), 3799-3800.-   72. Kwan, A. L. C. et al. (2005) “Evaluation of X-Ray Scatter    Properties in a Dedicated Cone-Beam Breast Ct Scanner,” Med. Phys.    32(9), 2967-2975.-   73. Graham, S. A. et al. (2007) “Compensators for Dose and Scatter    Management in Cone-Beam Computed Tomography,” Med. Phys. 34(7),    2691-2703.-   74. Mail, N. et al. (2009) “The Influence of Bowtie Filtration on    Cone-Beam Ct Image Quality,” Med. Phys. 36(1), 22-32.-   75. Menser, B. et al. (2010) “Use of Beam Shapers for Cone-Beam Ct    with Off-Centered Flat Detector,” pp 762233-762233-762212.-   76. Ning, R. et al. (2002) “X-Ray Scatter Suppression Algorithm for    Cone-Beam Volume Ct,” pp 774-781.-   77. Sun, M. et al. (2011) “Correction for Patient Table-Induced    Scattered Radiation in Cone-Beam Computed Tomography (Cbct),” Med.    Phys. 38(4), 2058-2073.-   78. Zhu, L. et al. (2006) “Scatter Correction Method for X-Ray Ct    Using Primary Modulation: Theory and Preliminary Results,” IEEE    Trans. Med. Imaging 25(12), 1573-1587.-   79. Zbijewski, W. and Beekman, F. J. (2006) “Efficient Monte Carlo    Based Scatter Artifact Reduction in Cone-Beam Micro-Ct,” IEEE Trans.    Med. Imaging 25(7), 817-827.-   80. Siewerdsen, J. H. et al. (2006) “A Simple, Direct Method for    X-Ray Scatter Estimation and Correction in Digital Radiography and    Cone-Beam Ct,” Med. Phys. 33(1), 187-197.-   81. Bertram, M. et al. (2006) “Scatter Correction for Cone-Beam    Computed Tomography Using Simulated Object Models,” pp    61421C-61421C-61412.-   82. Wiegert, J. et al. (2008) “Iterative Scatter Correction Based on    Artifact Assessment,” pp 69132B-69132B-69112.-   83. Poludniowski, G. et al. (2009) “An Efficient Monte Carlo-Based    Algorithm for Scatter Correction in Key Cone-Beam Ct,” Phys. Med.    Biol. 54(12), 3847-3864.-   84. Sechopoulos, I. (2012) “X-Ray Scatter Correction Method for    Dedicated Breast Computed Tomography,” Med. Phys. 39(5), 2896-2903.-   85. Landry, G. et al. (2015) “Investigating Ct to Cbct Image    Registration for Head and Neck Proton Therapy as a Tool for Daily    Dose Recalculation,” Med. Phys. 42(3), 1354-1366.-   86. Makarova, O. V. et al. (2002) “Focused Two-Dimensional    Antiscatter Grid for Mammography,” pp 148-155.-   87. Patel, T. et al. (2016) “Effects on Image Quality of a 2D    Antiscatter Grid in X-Ray Digital Breast Tomosynthesis: Initial    Experience Using the Dual Modality (X-Ray and Molecular)-   Breast Tomosynthesis Scanner,” Med. Phys. 43(4), 1720.-   88. Vogtmeier, G. et al. (2008) “Two-Dimensional Anti-Scatter Grids    for Computed Tomography Detectors,” Proc. SPIE-Int. Soc. Opt. Eng.    6913, 691359-691359-691311.-   89. Draper, N. R. and Smith, H. (2014) Applied Regression Analysis,    Third Edition ed., John Wiley & Sons.-   90. Lazos, D. and Williamson, J. F. (2012) “Impact of Flat    Panel-Imager Veiling Glare on Scatter-Estimation Accuracy and Image    Quality of a Commercial on-Board Cone-Beam Ct Imaging System,” Med.    Phys. 39(9), 5639-5651.-   91. Chan, H. P. and Doi, K. (1982) “Investigation of the Performance    of Antiscatter Grids: Monte Carlo Simulation Studies,” Phys. Med.    Biol. 27(6), 785-803.-   92. Chan, H. P. et al. (1985) “Performance of Antiscatter Grids in    Diagnostic Radiology: Experimental Measurements and Monte Carlo    Simulation Studies,” Med. Phys. 12(4), 449-454.-   93. Persliden, J. and Carlsson, G. A. (1997) “Scatter Rejection by    Air Gaps in Diagnostic Radiology. Calculations Using a Monte Carlo    Collision Density Method and Consideration of Molecular Interference    in Coherent Scattering,” Phys. Med. Biol. 42(1), 155-175.-   94. Bootsma, G. J. et al. (2011) “The Effects of Compensator and    Imaging Geometry on the Distribution of X-Ray Scatter in Cbct,” Med.    Phys. 38(2), 897-914.-   95. Wiegert, J. and Bertram, M. (2006) “Scattered Radiation in    Flat-Detector Based Cone-Beam Ct: Analysis of Voxelized Patient    Simulations,” pp 614235-614235-614212.-   96. Patel, T. et al. (2016) “Design and Evaluation of a Grid    Reciprocation Scheme for Use in Digital Breast Tomosynthesis,” pp    978805-978805-978819.-   97. Lin, C.-Y. et al. (2006) “A Study of Grid Artifacts Formation    and Elimination in Computed Radiographic Images,” J. Digit. Imaging    19(4), 351-361.-   98. Sasada, R. et al. (2003) “Stationary Grid Pattern Removal Using    2D Technique for Moire-Free Radiographic Image Display,” pp 688-697.-   99. Tang, H. et al. (2015) “A New Stationary Gridline Artifact    Suppression Method Based on the 2D Discrete Wavelet Transform,” Med.    Phys. 42(4), 1721-1729.-   100. Sun, X.-D. “X-Ray Detectors with a Grid Structured    Scintillators,” United States Patent Application Publication Number    US 2004-0251420 A1, application Ser. No. 10/866,408, filed Jun.    12, 2004. (published Dec. 16, 2004).-   101. Tang, C.-M. “Two-Dimensional, Anti-Scatter Grid and Collimator    Designs, and Its Motion, Fabrication and Assembly,” U.S. Pat. No.    6,252,938, application Ser. No. 09/459,597, filed Dec. 13, 1999.    (issued Jun. 26, 2001).-   102. Altunbas, C. et al. (2017) “Transmission Characteristics of a    Two Dimensional Antiscatter Grid Prototype for Cbct,” Med. Phys.    44(8), 3952-3964.

What is claimed is:
 1. A method of operating an x-ray imaging system,the method comprising: arranging an x-ray source and a detector havingan antiscatter grid (ASG) in a centered geometry wherein through-holesof the ASG are aligned with x-rays emission paths of the x-ray source;positioning an object between the detector and the x-ray source; imagingthe object in a first field of view (FOV) with the detector and thex-ray source arranged in the centered geometry; arranging the x-raysource and the detector in an offset geometry wherein the through-holesof the ASG are at least partially unaligned with the x-rays emissionpaths of the x-ray source; and moving the detector in the offsetgeometry to realign the through-holes of the ASG with the x-ray emissionpaths of the x-ray source.
 2. The method of claim 1 further comprisingimaging the object in at least a second FOV with the detector and thex-ray source arranged in the offset geometry.
 3. The method of claim 1,wherein the detector is a flat panel detector.
 4. The method of claim 1,wherein arranging the x-ray source and the detector in the offsetgeometry comprises shifting the detector laterally by a first distancex.
 5. The method of claim 4, wherein the moving step comprises tiltingthe detector upward toward the x-ray source by an angle (θ).
 6. Themethod of claim 5, wherein the angle (θ) is defined as:$\theta = {\arcsin\frac{x}{SDD}}$
 7. The method of claim 5, wherein themoving step further comprises upwardly shifting, by the tilting step, acenter of the detector by a second distance (y).
 8. The method of claim7, wherein the second distance (y) is defined as:y=SDD−√{square root over (SDD² −x ²)}
 9. A portable computed tomography(CT) system comprising: an O-shaped gantry defining an opening for ato-be-imaged object to be placed; an x-ray source operably coupled tothe O-shaped gantry; a flat panel detector (FPD) operably coupled to theO-shaped gantry, wherein the FPD comprises an x-ray absorbing sensorlayer and a detector pixel array; and an antiscatter grid (ASG) operablycoupled to a side of the FPD facing the opening of the O-shaped gantry,the ASG comprising a plurality of vertical walls defining open-endedchannels and formed of a radiation-absorbing material, the open-endedchannels arranged in a geometric pattern pointed toward the x-ray sourceto receive x-rays in an x-ray emission path from the x-ray source,wherein a pitch between the vertical walls of the ASG is larger than apitch of the detector pixel array.
 10. The portable CT scanner system ofclaim 9, wherein the ASG is a one-dimensional ASG.
 11. The portable CTscanner system of claim 9, wherein the ASG is a two-dimensional ASG. 12.The portable CT scanner system of claim 9, wherein a footprint of theASG introduces a pattern of signal intensity variations in projectionsof images of the object acquired by the detector pixel array.
 13. Theportable CT scanner system of claim 9, wherein the x-ray absorbingsensor layer is continuous, and wherein the ASG is placed directly onthe x-ray absorbing sensor layer.
 14. The portable CT scanner system ofclaim 9 further comprising a gap between the ASG and the x-ray absorbingsensor layer.
 15. The portable CT scanner system of claim 9, wherein theASG is on the FPD in the absence of the plurality of vertical wallsbeing aligned with pixels of the detector pixel array.
 16. The portableCT scanner system of claim 9, further comprising at least one controlleror computer operably coupled to the FPD, and configured to receive, fromthe FPD, data representative of an image of the object to facilitate anx-ray image of at least a portion of the object to be generated based onthe data representative of the image of the object.
 17. The portable CTscanner system of claim 16, wherein the at least one controller orcomputer is further configured to estimate a residual scatter intensityreaching the FPD.
 18. The portable CT scanner system of claim 17,wherein the at least one controller or computer is further configured tocorrect the residual scatter intensity by measuring signal intensityvariations in the FPD or the detector pixel array to facilitate animprovement in a quality of the x-ray image.
 19. An image datacorrection method for an x-ray imaging system, the method comprising:receiving, from a flat panel detector (FPD) of the x-ray imaging system,data representative of an x-ray image of at least a portion of anobject; and estimating a residual scatter intensity of x-ray radiationreaching the FPD from an x-ray source of the x-ray imaging system. 20.The method of claim 19, wherein the FPD include a two-dimensionalantiscatter grid (2D ASG) having a plurality of open-ended channels, themethod further comprising: determining a change in a ratio of an imageintensity underneath a footprint of the 2D ASG to an image intensity inopen-ended channels of 2D ASG in the absence of the object positionedbetween the x-ray source and the FPD; and determining a change in aratio of an image intensity underneath a footprint of the 2D ASG to animage intensity in open-ended channels of 2D ASG in the presence of theobject positioned between the x-ray source and the FPD.